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United States Patent Application |
20020111590
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Kind Code
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A1
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Davila, Luis A.
;   et al.
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August 15, 2002
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Medical devices, drug coatings and methods for maintaining the drug
coatings thereon
Abstract
Medical devices, and in particular implantable medical devices, may be
coated to minimize or substantially eliminate a biological organism's
reaction to the introduction of the medical device to the organism. The
medical devices may be coated with any number of biocompatible materials.
Therapeutic drugs, agents or compounds may be mixed with the
biocompatible materials and affixed to at least a portion of the medical
device. These therapeutic drugs, agents or compounds may also further
reduce a biological organism's reaction to the introduction of the
medical device to the organism. Various materials and coating
methodologies may be utilized to maintain the drugs, agents or compounds
on the medical device until delivered and positioned.
Inventors: |
Davila, Luis A.; (Pleasanton, CA)
; Lentz, David Christian; (Weston, FL)
; Llanos, Gerard H.; (Stewartsville, NJ)
; Mendez, Jorge Orlando; (Miami, FL)
; Narayanan, Pallassana V.; (Belle Mead, NJ)
; Pelton, Alan Roy; (Fremont, CA)
; Roller, Mark B.; (North Brunswick, NJ)
; Scheidt, Karl K.; (Pembroke Pines, FL)
; Scopelianos, Angelo George; (Whitehouse Station, NJ)
; Shaw, William Douglas JR.; (Miami, FL)
; Silver, James H.; (Redwood City, CA)
; Spaltro, John; (Asbury, NJ)
; Trepanier, Christine; (Fremont, CA)
; Wilson, David J.; (Ft. Lauderdale, FL)
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Correspondence Address:
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AUDLEY A. CIAMPORCERO JR.
JOHNSON & JOHNSON
ONE JOHNSON & JOHNSON PLAZA
NEW BRUNSWICK
NJ
08933-7003
US
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Serial No.:
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962496 |
Series Code:
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09
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Filed:
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September 25, 2001 |
Current U.S. Class: |
604/265; 623/1.15; 623/1.42 |
Class at Publication: |
604/265; 623/1.15; 623/1.42 |
International Class: |
A61M 005/32; A61M 025/00 |
Claims
What is claimed is:
1. A medical device for implantation into a treatment site of a living
organism, comprising: a biocompatible vehicle affixed to at least a
portion of the medical device; and at least one agent in therapeutic
dosages incorporated into the biocompatible vehicle for the treatment of
reactions by the living organism caused by the medical device or the
implantation thereof.
2. The medical device for implantation into a treatment site of a living
organism according to claim 1, wherein the biocompatible vehicle
comprises a polymeric matrix.
3. The medical device for implantation into a treatment site of a living
organism according to claim 2, wherein the polymeric matrix comprises
poly(ethylene-co-vinylacetate) and polybutylmethacrylate.
4. The medical device for implantation into a treatment site of a living
organism according to claim 2, wherein the polymeric matrix comprises
first and second layers, the first layer making contact with at least a
portion of the medical device and comprising a solution of
poly(ethylene-co-vinylacetate) and polybutylmethacrylate, and the second
layer comprising polybutylmethacrylate.
5. The medical device for implantation into a treatment site of a living
organism according to claim 4, wherein the at least one agent is
incorporated into the first layer.
6. The medical device for implantation into a treatment site of a living
organism according to claim 1, wherein the biocompatible vehicle
comprises a polyfluoro copolymer comprising polymerized residue of a
first moiety selected from the group consisting of vinylidenefluoride and
tetrafluoroethylene, and polymerized residue of a second moiety other
than the first moiety and which is copolymerized with the first moiety,
thereby producing the polyfluoro copolymer, wherein the relative amounts
of the polymerized residue of the first moiety and the polymerized
residue of the second moiety are effective to produce the biocompatible
coating with properties effective for use in coating implantable medical
devices when the coated medical device is subjected to a predetermined
maximum temperature, and a solvent in which the polyfluoro copolymer is
substantially soluble.
7. The medical device for implantation into a treatment site of a living
organism according to claim 6, wherein the polyfluoro copolymer comprises
from about 50 to about 92 weight percent of the polymerized residue of
the first moiety copolymerized with from about 50 to about 8 weight
percent of the polymerized residue of the second moiety.
8. The medical device for implantation into a treatment site of a living
organism according to claim 6, wherein said polyfluoro copolymer
comprises from about 50 to about 85 weight percent of polymerized residue
of vinylidenefluoride copolymerized with from about 50 to about 15 weight
percent of the polymerized residue of the second moiety.
9. The medical device for implantation into a treatment site of a living
organism according to claim 6, wherein said copolymer comprises from
about 55 to about 65 weight percent of the polymerized residue of the
vinylidenefluoride copolymerized with from about 45 to about 35 weight
percent of the polymerized residue of the second moiety.
10. The medical device for implantation into a treatment site of a living
organism according to claim 6, wherein the second moiety is selected from
the group consisting of hexafluoropropylene, tetrafluoroethylene,
vinylidenefluoride, 1-hydropentafluoropropylene, perfluoro (methyl vinyl
ether), chlorotrifluoroethylene, pentafluoropropene, trifluoroethylene,
hexafluoroacetone and hexafluoroisobutylene.
11. The medical device for implantation into a treatment site of a living
organism according to claim 6, wherein the second moiety is
hexafluoropropylene.
12. The medical device for implantation into a treatment site of a living
organism according to claim 1, wherein the at least one agent comprises
an anti-proliferative.
13. The medical device for implantation into a treatment site of a living
organism according to claim 1, wherein the at least one agent comprises
an anti-inflammatory.
14. The medical device for implantation into a treatment site of a living
organism according to claim 1, wherein the at least one agent comprises
an anti-coagulant.
15. The medical device for implantation into a treatment site of a living
organism according to claim 1, wherein the at least one agent comprises
rapamycin.
16. The medical device for implantation into a treatment site of a living
organism according to claim 1, wherein the at least one agent comprises
heparin.
17. A medical device for implantation into a treatment site of a living
organism, comprising: a biocompatible vehicle affixed to at least a
portion of the medical device; at least one agent in therapeutic dosages
incorporated into the biocompatible vehicle for the treatment of
reactions by the living organism caused by the medical device or the
implantation thereof; and a material for preventing the at least one
agent for separating from the medical device prior to and during
implantation of the medical device at the treatment site, the material
being affixed to at least one of the medical device or a delivery system
for the medical device.
18. The medical device for implantation into a treatment site of a living
organism according to claim 17, wherein the biocompatible vehicle
comprises a polymeric matrix.
19. The medical device for implantation into a treatment site of a living
organism according to claim 18, wherein the polymeric matrix comprises
poly(ethylene-co-vinylacetate) and polybutylmethacrylate.
20. The medical device for implantation into a treatment site of a living
organism according to claim 18, wherein the polymeric matrix comprises
first and second layers, the first layer making contact with at least a
portion of the medical device and comprising a solution of
poly(ethylene-co-vinylacetate) and polybutylmethacrylate, and the second
layer comprising polybutylmethacrylate.
21. The medical device for implantation into a treatment site of a living
organism according to claim 20, wherein the at least one agent is
incorporated into the first layer.
22. The medical device for implantation into a treatment site of a living
organism according to claim 17, wherein the biocompatible vehicle
comprises a polyfluoro copolymer comprising polymerized residue of a
first moiety selected from the group consisting of vinylidenefluoride and
tetrafluoroethylene, and polymerized residue of a second moiety other
than the first moiety and which is copolymerized with the first moiety,
thereby producing the polyfluoro copolymer, wherein the relative amounts
of the polymerized residue of the first moiety and the polymerized
residue of the second moiety are effective to produce the biocompatible
vehicle with properties effective for use in coating implantable medical
devices when the coated medical device is subjected to a predetermined
maximum temperature, and a solvent in which the polyfluoro copolymer is
substantially soluble.
23. The medical device for implantation into a treatment site of a living
organism according to claim 22, wherein the polyfluoro copolymer
comprises from about 50 to about 92 weight percent of the polymerized
residue of the first moiety copolymerized with from about 50 to about 8
weight percent of the polymerized residue of the second moiety.
24. The medical device for implantation into a treatment site of a living
organism according to claim 22, wherein said polyfluoro copolymer
comprises from about 50 to about 85 weight percent of polymerized residue
of vinylidenefluoride copolymerized with from about 50 to about 15 weight
percent of the polymerized residue of the second moiety.
25. The medical device for implantation into a treatment site of a living
organism according to claim 22, wherein said copolymer comprises from
about 55 to about 65 weight percent of the polymerized residue of the
vinylidenefluoride copolymerized with from about 45 to about 35 weight
percent of the polymerized residue of the second moiety.
26. The medical device for implantation into a treatment site of a living
organism according to claim 22, wherein the second moiety is selected
from the group consisting of hexafluoropropylene, tetrafluoroethylene,
vinylidenefluoride, 1-hydropentafluoropropylene, perfluoro (methyl vinyl
ether), chlorotrifluoroethylene, pentafluoropropene, trifluoroethylene,
hexafluoroacetone and hexafluoroisobutylene.
27. The medical device for implantation into a treatment site of a living
organism according to claim 22, wherein the second moiety is
hexafluoropropylene.
28. The medical device for implantation into a treatment site of a living
organism according to claim 17, wherein the material for preventing the
at least one agent from separating from the medical device comprises a
lubricious coating.
29. The medical device for implantation into a treatment site of a living
organism according to claim 28, wherein the lubricious coating comprises
a silicone-based material.
30. The medical device for implantation into a treatment site of a living
organism according to claim 28, wherein the lubricious coating is
incorporated into the medical device.
31. The medical device for implantation into a treatment site of a living
organism according to claim 28, wherein the lubricious coating is
incorporated into the delivery system for the medical device.
32. The medical device for implantation into a treatment site of a living
organism according to claim 17, wherein the material for preventing the
at least one agent from separating from the medical device comprises a
water soluble powder.
33. The medical device for implantation into a treatment site of a living
organism according to claim 32, wherein the water soluble powder is
incorporated onto the medical device.
34. The medical device for implantation into a treatment site of a living
organism according to claim 33, wherein the water soluble powder
comprises an anti-oxidant.
35. The medical device for implantation into a treatment site of a living
organism according to claim 33, wherein the water soluble powder
comprises an anti-coagulant.
36. A medical device for implantation into a treatment site of a living
organism, comprising: a stent; a biocompatible vehicle affixed to at
least a portion of the stent; and at least one agent in therapeutic
dosages incorporated into the biocompatible vehicle for the treatment of
reactions by the living organism caused by the medical device or the
implantation thereof.
37. The medical device for implantation into a treatment site of a living
organism according to claim 36, wherein the stent comprises a
substantially tubular member having open ends, and a first diameter for
insertion into a lumen of a vessel and a second diameter for anchoring in
the lumen of the vessel.
38. The medical device for implantation into a treatment site of a living
organism according to claim 37, wherein the tubular member comprises a
plurality of adjacent hoops formed from a plurality of longitudinal
struts and a plurality of loops connecting adjacent struts, the adjacent
struts are connected at opposite ends to form a substantially S-shaped
pattern, and a plurality of bridges which connect adjacent hoops.
39. The medical device for implantation into a treatment site of a living
organism according to claim 38, wherein the biocompatible vehicle
comprises a polymeric matrix.
40. The medical device for implantation into a treatment site of a living
organism according to claim 39, wherein the polymeric matrix comprises
poly(ethylene-co-vinylacetate) and polybutylmethacrylate.
41. The medical device for implantation into a treatment site of a living
organism according to claim 39, wherein the polymeric matrix comprises
first and second layers, the first layer making contact with at least a
portion of the medical device and comprising a solution of
poly(ethylene-co-vinylacetate) and polybutylmethacrylate, and the second
layer comprising polybutylmethacrylate.
42. The medical device for implantation into a treatment site of a living
organism according to claim 41, wherein the at least one agent is
incorporated into the first layer.
43. The medical device for implantation into a treatment site of a living
organism according to claim 38, wherein the biocompatible vehicle
comprises a polyfluoro copolymer copolymer comprising polymerized residue
of a first moiety selected from the group consisting of
vinylidenefluoride and tetrafluoroethylene, and polymerized residue of a
second moiety other than the first moiety and which is copolymerized with
the first moiety, thereby producing the polyfluoro copolymer, wherein the
relative amounts of the polymerized residue of the first moiety and the
polymerized residue of the second moiety are effective to produce the
biocompatible coating with properties effective for use in coating
implantable medical devices when the coated medical device is subjected
to a predetermined maximum temperature, and a solvent in which the
polyfluoro copolymer is substantially soluble.
44. The medical device for implantation into a treatment site of a living
organism according to claim 43, wherein the polyfluoro copolymer
comprises from about 50 to about 92 weight percent of the polymerized
residue of the first moiety copolymerized with from about 50 to about 8
weight percent of the polymerized residue of the second moiety.
45. The medical device for implantation into a treatment site of a living
organism according to claim 43, wherein said polyfluoro copolymer
comprises from about 50 to about 85 weight percent of polymerized residue
of vinylidenefluoride copolymerized with from about 50 to about 15 weight
percent of the polymerized residue of the second moiety.
46. The medical device for implantation into a treatment site of a living
organism according to claim 43, wherein said copolymer comprises from
about 55 to about 65 weight percent of the polymerized residue of the
vinylidenefluoride copolymerized with from about 45 to about 35 weight
percent of the polymerized residue of the second moiety.
47. The medical device for implantation into a treatment site of a living
organism according to claim 43, wherein the second moiety is selected
from the group consisting of hexafluoropropylene, tetrafluoroethylene,
vinylidenefluoride, 1-hydropentafluoropropylene, perfluoro (methyl vinyl
ether), chlorotrifluoroethylene, pentafluoropropene, trifluoroethylene,
hexafluoroacetone and hexafluoroisobutylene.
48. The medical device for implantation into a treatment site of a living
organism according to claim 43, wherein the second moiety is
hexafluoropropylene.
49. The medical device for implantation into a treatment site of a living
organism according to claim 38, wherein the at least one agent comprises
an anti-proliferative.
50. The medical device for implantation into a treatment site of a living
organism according to claim 38, wherein the at least one agent comprises
an anti-inflammatory.
51. The medical device for implantation into a treatment site of a living
organism according to claim 38, wherein the at least one agent comprises
an anti-coagulant.
52. The medical device for implantation into a treatment site of a living
organism according to claim 38, wherein the at least one agent comprises
rapamycin.
53. The medical device for implantation into a treatment site of a living
organism according to claim 38, wherein the at least one agent comprises
heparin.
54. A medical device for implantation into a treatment site of a living
organism, comprising: a stent having a substantially tubular member
having open ends, and a first diameter for insertion into a lumen of a
vessel and a second diameter for anchoring in the lumen of the vessel; a
biocompatible vehicle affixed to at least a portion of the stent; at
least one agent in therapeutic dosages incorporated into the
biocompatible vehicle for the treatment of reactions by the living
organism caused by the medical device or the implantation thereof; and a
material for preventing the at least one agent from separating from the
medical device prior to and during implantation of the medical device at
the treatment site, the material being affixed to at least one of the
medical devices or a delivery system for the medical device.
55. The medical device for implantation into a treatment site of a living
organism according to claim 54, wherein the tubular member comprises a
plurality of adjacent hoops formed from a plurality of longitudinal
struts and a plurality of loops connecting adjacent struts, the adjacent
struts are connected at opposite ends to form a substantially S-shaped
pattern, and a plurality of bridges which connect adjacent hoops.
56. The medical device for implantation into a treatment site of a living
organism according to claim 55, wherein the stent comprises a
superelastic alloy.
57. The medical device for implantation into a treatment site of a living
organism according to claim 56, wherein the superelastic alloy comprises
from about fifty percent to about sixty percent Nickel and the remainder
Titanium.
58. The medical device for implantation into a treatment site of a living
organism according to claim 57, wherein the biocompatible coating
comprises a polymeric matrix.
59. The medical device for implantation into a treatment site of a living
organism according to claim 58, wherein the polymeric matrix comprises
poly(ethylene-co-vinylacetate) and polybutylmethacrylate.
60. The medical device for implantation into a treatment site of a living
organism according to claim 58, wherein the polymeric matrix comprises
first and second layers, the first layer making contact with at least a
portion of the medical device and comprising a solution of
poly(ethylene-co-vinylacetate) and polybutylmethacrylate, and the second
layer comprising polybutylmethacrylate.
61. The medical device for implantation into a treatment site of a living
organism according to claim 60, wherein the at least one agent is
incorporated into the first layer.
62. The medical device for implantation into a treatment site of a living
organism according to claim 57, wherein the biocompatible vehicle
comprises a polyfluoro copolymer comprising polymerized residue of a
first moiety selected from the group consisting of vinylidenefluoride and
tetrafluoroethylene, and polymerized residue of a second moiety other
than the first moiety and which is copolymerized with the first moiety,
thereby producing the polyfluoro copolymer, wherein the relative amounts
of the polymerized residue of the first moiety and the polymerized
residue of the second moiety are effective to produce the biocompatible
coating with properties effective for use in coating implantable medical
devices when the coated medical device is subjected to a predetermined
maximum temperature, and a solvent in which the polyfluoro copolymer is
substantially soluble.
63. The medical device for implantation into a treatment site of a living
organism according to claim 62, wherein the polyfluoro copolymer
comprises from about 50 to about 92 weight percent of the polymerized
residue of the first moiety copolymerized with from about 50 to about 8
weight percent of the polymerized residue of the second moiety.
64. The medical device for implantation into a treatment site of a living
organism according to claim 62, wherein said polyfluoro copolymer
comprises from about 50 to about 85 weight percent of polymerized residue
of vinylidenefluoride copolymerized with from about 50 to about 15 weight
percent of the polymerized residue of the second moiety.
65. The medical device for implantation into a treatment site of a living
organism according to claim 62, wherein said copolymer comprises from
about 55 to about 65 weight percent of the polymerized residue of the
vinylidenefluoride copolymerized with from about 45 to about 35 weight
percent of the polymerized residue of the second moiety.
66. The medical device for implantation into a treatment site of a living
organism according to claim 62, wherein the second moiety is selected
from the group consisting of hexafluoropropylene, tetrafluoroethylene,
vinylidenefluoride, 1-hydropentafluoropropylene, perfluoro (methyl vinyl
ether), chlorotrifluoroethylene, pentafluoropropene, trifluoroethylene,
hexafluoroacetone and hexafluoroisobutylene.
67. The medical device for implantation into a treatment site of a living
organism according to claim 62, wherein the second moiety is
hexafluoropropylene.
68. The medical device for implantation into a treatment site of a living
organism according to claim 55, wherein the at least one agent comprises
an anti-proliferative.
69. The medical device for implantation into a treatment site of a living
organism according to claim 55, wherein the at least one agent comprises
an anti-inflammatory.
70. The medical device for implantation into a treatment site of a living
organism according to claim 55, wherein the at least one agent comprises
an anti-coagulant.
71. The medical device for implantation into a treatment site of a living
organism according to claim 55, wherein the at least one agent comprises
rapamycin.
72. The medical device for implantation into a treatment site of a living
organism according to claim 55, wherein the at least one agent comprises
heparin.
73. The medical device for implantation into a treatment site of a living
organism according to claim 55, wherein the material for preventing the
at least one agent from separating from the medical device comprises a
lubricious coating.
74. The medical device for implantation into a treatment site of a living
organism according to claim 73, wherein the lubricious coating comprises
a silicone-based material.
75. The medical device for implantation into a treatment site of a living
organism according to claim 73, wherein the lubricious coating is
incorporated into the medical device.
76. The medical device for implantation into a treatment site of a living
organism according to claim 73, wherein the lubricious coating is
incorporated into the delivery system for the medical device.
77. The medical device for implantation into a treatment site of a living
organism according to claim 55, wherein the material for preventing the
at least one agent from separating from the medical device comprises a
water soluble powder.
78. The medical device for implantation into a treatment site of a living
organism according to claim 77, wherein the water soluble powder is
incorporated onto the medical device.
79. The medical device for implantation into a treatment site of a living
organism according to claim 78, wherein the water soluble powder
comprises an anti-oxidant.
80. The medical device for implantation into a treatment site of a living
organism according to claim 78, wherein the water soluble powder
comprises an anti-coagulant.
81. The medical device for implantation into a treatment site of a living
organism according to claim 75, wherein the lubricious coating is
incorporated into the polymeric matrix.
82. The medical device for implantation into a treatment site of a living
organism according to claim 78, wherein the water soluble powder is
affixed to the surface of the polymeric matrix.
83. The medical device for implantation into a treatment site of a living
organism according to claim 57, further comprises at least one marker
connected to at least one end of the substantially tubular member, the at
least one marker comprising a marker housing and a marker insert having a
radius of curvature equal to the radius of curvature of the substantially
tubular member.
84. The medical device for implantation into a treatment site of a living
organism according to claim 83, wherein the marker housing comprises the
same material as the stent and is integral thereto, thereby forming a
unitary structure.
85. The medical device for implantation into a treatment site of a living
organism according to claim 84, wherein the marker insert comprises a
material having a radiopacity higher than that of the material comprising
the stent.
86. The medical device for implantation into a treatment site of a living
organism according to claim 85, wherein the marker insert comprises
Tantalum.
87. The medical device for implantation into a treatment site of a living
organism according to claim 86, wherein the marker insert is secured in
the marker housing by frictional, locking engagement.
88. The medical device for implantation into a treatment site of a living
organism according to claim 87, wherein the marker insert is secured in
the marker housing by a protruding ridge.
89. A local drug delivery device comprising: a stent having a
substantially tubular member having open ends, and a first diameter for
insertion into a lumen of a vessel and a second diameter for anchoring in
the lumen of a vessel; a biocompatible polymeric vehicle affixed to at
least a portion of the stent; and rapamycin, in therapeutic dosages,
incorporated into the polymeric vehicle.
90. The local drug delivery device according to claim 89, wherein from
about fifty micrograms to about one thousand micrograms of rapamycin is
provided per square centimeter of the lumen into which the stent is
anchored.
91. The local drug delivery device according to claim 89, wherein the
polymeric coating comprises a combination of vinylidene fluoride and
hexafluoropropylene.
92. The local drug delivery device according to claim 91, wherein the
polymeric coating weight is in the range from about two hundred to about
one thousand seven hundred micrograms per square centimeter of the lumen
into which the stent is anchored.
93. A method of coating a medical device with a therapeutic agent
comprising the steps of: creating a polymer utilizing vinylidene fluoride
and hexafluoropropylene in a batch emulsion polymerization process;
priming the medical device with the polymer utilizing a dip coating
process; creating a polymer and therapeutic agent mixture; applying the
polymer and therapeutic agent mixture on the primer layer utilizing a
spin coating process; and drying the medical device in a vacuum oven for
approximately sixteen hours at a temperature in the range of fifty to
sixty degrees centigrade.
94. A medical device for implantation into a treatment site of a living
organism, comprising: a biocompatible vehicle affixed to at least a
portion of the medical device; and at least one agent incorporated into
the biocompatible vehicle, the at least one agent being designed to react
with one or more chemicals produced by the living organism to treat
reactions by the living organism caused by the medical device or the
implantation thereof.
95. A medical device for implantation into the vasculature of a living
organism, comprising: a self-expanding stent; a biocompatible vehicle
affixed to at least a portion of the stent; and rapamycin, in therapeutic
dosages, incorporated into the biocompatible vehicle for the prevention
of restenosis.
96. A method of coating a medical device with a therapeutic agent
comprising the steps of: creating a polymer utilizing vinylidene fluoride
and hexafluoropropylene; adding one or more therapeutic agents to the
polymer to create a polymer and therapeutic agent mixture; and applying
the polymer and therapeutic agent mixture to the medical device.
97. A medical device for implantation into a treatment site of a living
organism, comprising: a biocompatible vehicle affixed to at least a
portion of the medical device; at least one agent in therapeutic dosages
incorporated into the biocompatible vehicle for the treatment of disease
proximate the implantation site.
98. A medical device for implantation into a treatment site of a living
organism, comprising: a biocompatible vehicle affixed to at least a
portion of the medical device; at least one agent in therapeutic dosages
incorporated into the biocompatible vehicle for the treatment of disease
remote from the implantation site.
99. A method of coating a medical device with a therapeutic agent
comprising the steps of: creating a polymer utilizing vinylidene fluoride
and hexafluoropropylene in a batch dispersion polymerization process;
priming the medical device with the polymer utilizing a dip coating
process; creating a polymer and therapeutic agent mixture; applying the
polymer and therapeutic agent mixture on the primer layer utilizing a
spin coating process; and drying the medical device in a vacuum oven for
approximately sixteen hours at a temperature in the range of fifty to
sixty degrees centigrade.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a continuation-in-part application of U.S.
application Ser. No. 09/887,464 filed Jun. 22, 2001, a
continuation-in-part application of U.S. application Ser. No. 09/675,882,
filed Sep. 29, 2000,a continuation-in-part application of U.S.
application Ser. No. 09/884,729 filed Jun. 19, 2001 and a
continuation-in-part of U.S. application Ser. No. 09/850,482 filed May 7,
2001.
BACKGROUND OF THE INVENTION
[0002] 1. Field of the Invention
[0003] The present invention relates to the local administration of
drug/drug combinations for the prevention and treatment of vascular
disease, and more particularly to intraluminal medical devices for the
local delivery of drug/drug combinations for the prevention and treatment
of vascular disease caused by injury and methods for maintaining the
drug/drug combinations on the intraluminal medical devices. The present
invention also relates to medical devices having drugs, agents or
compounds affixed thereto to minimize or substantially eliminate a
biological organism's reaction to the introduction of the medical device
to the organism.
[0004] 2. Discussion of the Related Art
[0005] Many individuals suffer from circulatory disease caused by a
progressive blockage of the blood vessels that perfuse the heart and
other major organs with nutrients. More severe blockage of blood vessels
in such individuals often leads to hypertension, ischemic injury, stroke,
or myocardial infarction. Atherosclerotic lesions, which limit or
obstruct coronary blood flow, are the major cause of ischemic heart
disease. Percutaneous transluminal coronary angioplasty is a medical
procedure whose purpose is to increase blood flow through an artery.
Percutaneous transluminal coronary angioplasty is the predominant
treatment for coronary vessel stenosis. The increasing use of this
procedure is attributable to its relatively high success rate and its
minimal invasiveness compared with coronary bypass surgery. A limitation
associated with percutaneous transluminal coronary angioplasty is the
abrupt closure of the vessel which may occur immediately after the
procedure and restenosis which occurs gradually following the procedure.
Additionally, restenosis is a chronic problem in patients who have
undergone saphenous vein bypass grafting. The mechanism of acute
occlusion appears to involve several factors and may result from vascular
recoil with resultant closure of the artery and/or deposition of blood
platelets and fibrin along the damaged length of the newly opened blood
vessel.
[0006] Restenosis after percutaneous transluminal coronary angioplasty is
a more gradual process initiated by vascular injury. Multiple processes,
including thrombosis, inflammation, growth factor and cytokine release,
cell proliferation, cell migration and extracellular matrix synthesis
each contribute to the restenotic process.
[0007] While the exact mechanism of restenosis is not completely
understood, the general aspects of the restenosis process have been
identified. In the normal arterial wall, smooth muscle cells proliferate
at a low rate, approximately less than 0.1 percent per day. Smooth muscle
cells in the vessel walls exist in a contractile phenotype characterized
by eighty to ninety percent of the cell cytoplasmic volume occupied with
the contractile apparatus. Endoplasmic reticulum, Golgi, and free
ribosomes are few and are located in the perinuclear region.
Extracellular matrix surrounds the smooth muscle cells and is rich in
heparin-like glycosylaminoglycans which are believed to be responsible
for maintaining smooth muscle cells in the contractile phenotypic state
(Campbell and Campbell, 1985).
[0008] Upon pressure expansion of an intracoronary balloon catheter during
angioplasty, smooth muscle cells within the vessel wall become injured,
initiating a thrombotic and inflammatory response. Cell derived growth
factors such as platelet derived growth factor, basic fibroblast growth
factor, epidermal growth factor, thrombin, etc., released from platelets,
invading macrophages and/or leukocytes, or directly from the smooth
muscle cells provoke a proliferative and migratory response in medial
smooth muscle cells. These cells undergo a change from the contractile
phenotype to a synthetic phenotype characterized by only a few
contractile filament bundles, extensive rough endoplasmic reticulum,
Golgi and free ribosomes. Proliferation/migration usually begins within
one to two days post-injury and peaks several days thereafter (Campbell
and Campbell, 1987; Clowes and Schwartz, 1985).
[0009] Daughter cells migrate to the intimal layer of arterial smooth
muscle and continue to proliferate and secrete significant amounts of
extracellular matrix proteins. Proliferation, migration and extracellular
matrix synthesis continue until the damaged endothelial layer is repaired
at which time proliferation slows within the intima, usually within seven
to fourteen days post-injury. The newly formed tissue is called
neointima. The further vascular narrowing that occurs over the next three
to six months is due primarily to negative or constrictive remodeling.
[0010] Simultaneous with local proliferation and migration, inflammatory
cells adhere to the site of vascular injury. Within three to seven days
post-injury, inflammatory cells have migrated to the deeper layers of the
vessel wall. In animal models employing either balloon injury or stent
implantation, inflammatory cells may persist at the site of vascular
injury for at least thirty days (Tanaka et al., 1993; Edelman et al.,
1998). Inflammatory cells therefore are present and may contribute to
both the acute and chronic phases of restenosis.
[0011] Numerous agents have been examined for presumed anti-proliferative
actions in restenosis and have shown some activity in experimental animal
models. Some of the agents which have been shown to successfully reduce
the extent of intimal hyperplasia in animal models include: heparin and
heparin fragments (Clowes, A. W. and Karnovsky M., Nature 265: 25-26,
1977; Guyton, J. R. et al., Circ. Res., 46: 625-634, 1980; Clowes, A. W.
and Clowes, M. M., Lab. Invest. 52: 611-616,1985; Clowes, A. W. and
Clowes, M. M., Circ. Res. 58: 839-845, 1986; Majesky et al., Circ. Res.
61: 296-300, 1987; Snow et al., Am. J. Pathol. 137: 313-330, 1990; Okada,
T. et al., Neurosurgery 25: 92-98, 1989), colchicine (Currier, J. W. et
al., Circ. 80: 11-66,1989), taxol (Sollot, S. J. et al., J. Clin. Invest.
95: 1869-1876, 1995), angiotensin converting enzyme (ACE) inhibitors
(Powell, J. S. et al., Science, 245: 186-188,1989), angiopeptin
(Lundergan, C. F. et al. Am. J. Cardiol. 17(Suppl. B):132B-136B, 1991),
cyclosporin A (Jonasson, L. et al., Proc. Natl., Acad. Sci., 85:
2303,1988), goat-anti-rabbit PDGF antibody (Ferns, G. A. A., et al.,
Science 253: 1129-1132, 1991), terbinafine (Nemecek, G. M. et al., J.
Pharmacol. Exp. Thera. 248: 1167-1174, 1989), trapidil (Liu, M. W. et
al., Circ. 81: 1089-1093,1990), tranilast (Fukuyama, J. et al., Eur. J.
Pharmacol. 318: 327-332, 1996), interferon-gamma (Hansson, G. K. and
Holm, J., Circ. 84:1266-1272,1991), rapamycin (Marx, S. O. et al., Circ.
Res. 76: 412-417, 1995), steroids (Colburn, M. D. et al., J. Vasc. Surg.
15: 510-518, 1992), see also Berk, B. C. et al., J. Am. Coll. Cardiol.
17: 111 B-117B, 1991), ionizing radiation (Weinberger, J. et al., Int. J.
Rad. Onc. Biol. Phys. 36: 767-775, 1996), fusion toxins (Farb, A. et al.,
Circ. Res. 80: 542-550,1997) antisense oligionucleotides (Simons, M. et
al., Nature 359: 67-70, 1992) and gene vectors (Chang, M. W. et al., J.
Clin. Invest. 96: 2260-2268, 1995). Anti-proliferative action on smooth
muscle cells in vitro has been demonstrated for many of these agents,
including heparin and heparin conjugates, taxol, tranilast, colchicine,
ACE inhibitors, fusion toxins, antisense oligionucleotides, rapamycin and
ionizing radiation. Thus, agents with diverse mechanisms of smooth muscle
cell inhibition may have therapeutic utility in reducing intimal
hyperplasia.
[0012] However, in contrast to animal models, attempts in human
angioplasty patients to prevent restenosis by systemic pharmacologic
means have thus far been unsuccessful. Neither aspirin-dipyridamole,
ticlopidine, anti-coagulant therapy (acute heparin, chronic warfarin,
hirudin or hirulog), thromboxane receptor antagonism nor steroids have
been effective in preventing restenosis, although platelet inhibitors
have been effective in preventing acute reocclusion after angioplasty
(Mak and Topol, 1997; Lang et al., 1991; Popma et al., 1991). The
platelet GP II.sub.bII.sub.a receptor, antagonist, Reopro.RTM. is still
under study but Reopro.RTM. has not shown definitive results for the
reduction in restenosis following angioplasty and stenting. Other agents,
which have also been unsuccessful in the prevention of restenosis,
include the calcium channel antagonists, prostacyclin mimetics,
angiotensin converting enzyme inhibitors, serotonin receptor antagonists,
and anti-proliferative agents. These agents must be given systemically,
however, and attainment of a therapeutically effective dose may not be
possible; anti-proliferative (or anti-restenosis) concentrations may
exceed the known toxic concentrations of these agents so that levels
sufficient to produce smooth muscle inhibition may not be reached (Mak
and Topol, 1997; Lang et al., 1991; Popma et al., 1991).
[0013] Additional clinical trials in which the effectiveness for
preventing restenosis utilizing dietary fish oil supplements or
cholesterol lowering agents has been examined showing either conflicting
or negative results so that no pharmacological agents are as yet
clinically available to prevent post-angioplasty restenosis (Mak and
Topol, 1997; Franklin and Faxon, 1993: Serruys, P. W. et al., 1993).
Recent observations suggest that the antilipid/antioxident agent,
probucol, may be useful in preventing restenosis but this work requires
confirmation (Tardif et al., 1997; Yokoi, et al., 1997). Probucol is
presently not approved for use in the United States and a thirty-day
pretreatment period would preclude its use in emergency angioplasty.
Additionally, the application of ionizing radiation has shown significant
promise in reducing or preventing restenosis after angioplasty in
patients with stents (Teirstein et al., 1997). Currently, however, the
most effective treatments for restenosis are repeat angioplasty,
atherectomy or coronary artery bypass grafting, because no therapeutic
agents currently have Food and Drug Administration approval for use for
the prevention of post-angioplasty restenosis.
[0014] Unlike systemic pharmacologic therapy, stents have proven useful in
significantly reducing restenosis. Typically, stents are
balloon-expandable slotted metal tubes (usually, but not limited to,
stainless steel), which, when expanded within the lumen of an
angioplastied coronary artery, provide structural support through rigid
scaffolding to the arterial wall. This support is helpful in maintaining
vessel lumen patency. In two randomized clinical trials, stents increased
angiographic success after percutaneous transluminal coronary
angioplasty, by increasing minimal lumen diameter and reducing, but not
eliminating, the incidence of restenosis at six months (Serruys et al.,
1994; Fischman et al., 1994).
[0015] Additionally, the heparin coating of stents appears to have the
added benefit of producing a reduction in sub-acute thrombosis after
stent implantation (Serruys et al., 1996). Thus, sustained mechanical
expansion of a stenosed coronary artery with a stent has been shown to
provide some measure of restenosis prevention, and the coating of stents
with heparin has demonstrated both the feasibility and the clinical
usefulness of delivering drugs locally, at the site of injured tissue.
[0016] As stated above, the use of heparin coated stents demonstrates the
feasibility and clinical usefulness of local drug delivery; however, the
manner in which the particular drug or drug combination is affixed to the
local delivery device will play a role in the efficacy of this type of
treatment. For example, the processes and materials utilized to affix the
drug/drug combinations to the local delivery device should not interfere
with the operations of the drug/drug combinations. In addition, the
processes and materials utilized should be biocompatible and maintain the
drug/drug combinations on the local device through delivery and over a
given period of time. For example, removal of the drug/drug combination
during delivery of the local delivery device may potentially cause
failure of the device.
[0017] Accordingly, there exists a need for drug/drug combinations and
associated local delivery devices for the prevention and treatment of
vascular injury causing intimal thickening which is either biologically
induced, for example atherosclerosis, or mechanically induced, for
example, through percutaneous transluminal coronary angioplasty. In
addition, there exists a need for maintaining the drug/drug combinations
on the local delivery device through delivery and positioning as well as
ensuring that the drug/drug combination is released in therapeutic
dosages over a given period of time.
[0018] A variety of stent coatings and compositions have been proposed for
the prevention and treatment of injury causing intimal thickening. The
coatings may be capable themselves of reducing the stimulus the stent
provides to the injured lumen wall, thus reducing the tendency towards
thrombosis or restenosis. Alternately, the coating may deliver a
pharmaceutical/therapeutic agent or drug to the lumen that reduces smooth
muscle tissue proliferation or restenosis. The mechanism for delivery of
the agent is through diffusion of the agent through either a bulk polymer
or through pores that are created in the polymer structure, or by erosion
of a biodegradable coating.
[0019] Both bioabsorbable and biostable compositions have been reported as
coatings for stents. They generally have been polymeric coatings that
either encapsulate a pharmaceutical/therapeutic agent or drug, e.g.
rapamycin, taxol etc., or bind such an agent to the surface, e.g.
heparin-coated stents. These coatings are applied to the stent in a
number of ways, including, though not limited to, dip, spray, or spin
coating processes.
[0020] One class of biostable materials that has been reported as coatings
for stents is polyfluoro homopolymers. Polytetrafluoroethylene (PTFE)
homopolymers have been used as implants for many years. These
homopolymers are not soluble in any solvent at reasonable temperatures
and therefore are difficult to coat onto small medical devices while
maintaining important features of the devices (e.g. slots in stents).
[0021] Stents with coatings made from polyvinylidenefluoride homopolymers
and containing pharmaceutical/therapeutic agents or drugs for release
have been suggested. However, like most crystalline polyfluoro
homopolymers, they are difficult to apply as high quality films onto
surfaces without subjecting them to relatively high temperatures, that
correspond to the melting temperature of the polymer.
[0022] It would be advantageous to develop coatings for implantable
medical devices that will reduce thrombosis, restenosis, or other adverse
reactions, that may include, but do not require, the use of
pharmaceutical or therapeutic agents or drugs to achieve such affects,
and that possess physical and mechanical properties effective for use in
such devices even when such coated devices are subjected to relatively
low maximum temperatures.
SUMMARY OF THE INVENTION
[0023] The drug/drug combination therapies, drug/drug combination carriers
and associated local delivery devices of the present invention provide a
means for overcoming the difficulties associated with the methods and
devices currently in use, as briefly described above. In addition, the
methods for maintaining the drug/drug combinations and drug/drug
combination carriers on the local delivery device ensure that the
drug/drug combination therapies reach the target site.
[0024] In accordance with one aspect, the present invention is directed to
a medical device for implantation into a treatment site of a living
organism. The device comprises a biocompatible vehicle affixed to at
least a portion of the medical device, and at least one agent in
therapeutic dosages incorporated into the biocompatible vehicle for the
treatment of reactions by the living organism caused by the medical
device or the implantation thereof.
[0025] In accordance with another aspect, the present invention is
directed to a medical device for implantation into a treatment site of a
living organism. The device comprises a biocompatible vehicle affixed to
at least a portion of the medical device, at least one agent in
therapeutic dosages incorporated into the biocompatible vehicle for the
treatment of reactions by the living organism caused by the medical
device or the implantation thereof, and a material for preventing the at
least one agent from separating from the medical device prior to and
during implantation of the medical device at the treatment site, the
material being affixed to at least one of the medical devices or a
delivery system for the medical device.
[0026] In accordance with another aspect, the present invention is
directed to a medical device for implantation into a treatment site of a
living organism. The device comprises a stent, a biocompatible vehicle
affixed to at least a portion of the stent, and at least one agent in
therapeutic dosages incorporated into the biocompatible vehicle for the
treatment of reactions by the living organism caused by the medical
device or the implantation thereof.
[0027] In accordance with another aspect, the present invention is
directed to a medical device for implantation into a treatment site of a
living organism. The device comprises a stent having a substantially
tubular member having open ends, and a first diameter for insertion into
a lumen of a vessel and a second diameter for anchoring in the lumen of
the vessel, a biocompatible vehicle affixed to at least a portion of the
stent, at least one agent in therapeutic dosages incorporated into the
biocompatible vehicle for the treatment of reactions by the living
organism caused by the medical device or the implantation thereof, and a
material for preventing the at least one agent from separating from the
medical device prior to and during implantation of the medical device at
the treatment site, the material being affixed to at least one of the
medical devices or a delivery system for the medical device.
[0028] In accordance with another aspect, the present invention is
directed to a local drug delivery device. The device comprises a stent
having a substantially tubular member having open ends, and a first
diameter for insertion into a lumen of a vessel and a second diameter for
anchoring in the lumen of a vessel, a biocompatible polymeric vehicle
affixed to at least a portion of the stent, and rapamycin, in therapeutic
dosages, incorporated into the biocompatible polymeric vehicle.
[0029] In accordance with another aspect, the present invention is
directed to a method of coating a medical device with a therapeutic
agent. The method comprises the steps of creating a polymer utilizing
vinylidene fluoride and hexafluoropropylene in a batch emulsion
polymerization process, priming the medical device with the polymer
utilizing a dip coating process, creating a polymer and therapeutic agent
mixture, applying the polymer and therapeutic agent mixture on the primer
layer utilizing a spin coating process, and drying the medical device in
a vacuum oven for approximately sixteen hours at a temperature in the
range of fifty to sixty degrees centigrade.
[0030] In accordance with another aspect, the present invention is
directed to a medical device for implantation into a treatment site of a
living organism. The medical device comprises a biocompatible vehicle
affixed to at least a portion of the medical device, and at least one
agent incorporated into the biocompatible vehicle. The at least one agent
being designed to react with one or more chemicals produced by the living
organism to treat reactions by the living organism caused by the medical
device or the implantation thereof.
[0031] In accordance with another aspect, the present invention is
directed to a medical device for implantation into the vasculature of a
living organism. The medical device comprises a self-expanding stent, a
biocompatible vehicle affixed to at least a portion of the stent, and
rapamycin, in therapeutic dosages, incorporated into the biocompatible
vehicle for the prevention of restenosis.
[0032] In accordance with another aspect, the present invention is
directed to a method of coating a medical device with a therapeutic
agent. The method comprises the steps of creating a polymer utilizing
vinylidene fluoride and hexafluoropropylene, adding one or more
therapeutic agents to the polymer to create a polymer and therapeutic
agent mixture, and applying the polymer and therapeutic agent mixture to
the medical device.
[0033] In accordance with another aspect, the present invention is
directed to a medical device for implantation into a treatment site of a
living organism. The medical device comprises a biocompatible vehicle
affixed to at least a portion of the medical device, at least one agent
in therapeutic dosages incorporated into the biocompatible vehicle for
the treatment of disease proximate the implantation site.
[0034] In accordance with another aspect, the present invention is
directed to a medical device for implantation into a treatment site of a
living organism. The medical device comprises a biocompatible vehicle
affixed to at least a portion of the medical device, at least one agent
in therapeutic dosages incorporated into the biocompatible vehicle for
the treatment of disease remote from the implantation site.
[0035] The medical devices, drug coatings and methods for maintaining the
drug coatings or vehicles thereon of the present invention utilizes a
combination of materials to treat disease, and reactions by living
organisms due to the implantation of medical devices for the treatment of
disease or other conditions. The local delivery of drugs, agents or
compounds generally substantially reduces the potential toxicity of the
drugs, agents or compounds when compared to systemic delivery while
increasing their efficacy.
[0036] Drugs, agents or compounds may be affixed to any number of medical
devices to treat various diseases. The drugs, agents or compounds may
also be affixed to minimize or substantially eliminate the biological
organism's reaction to the introduction of the medical device utilized to
treat a separate condition. For example, stents may be introduced to open
coronary arteries or other body lumens such as biliary ducts. The
introduction of these stents cause a smooth muscle cell proliferation
effect as well as inflammation. Accordingly, the stents may be coated
with drugs, agents or compounds to combat these reactions.
[0037] The drugs, agents or compounds will vary depending upon the type of
medical device, the reaction to the introduction of the medical device
and/or the disease sought to be treated. The type of coating or vehicle
utilized to immobilize the drugs, agents or compounds to the medical
device may also vary depending on a number of factors, including the type
of medical device, the type of drug, agent or compound and the rate of
release thereof.
[0038] In order to be effective, the drugs, agents or compounds should
preferably remain on the medical devices during delivery and
implantation. Accordingly, various coating techniques for creating strong
bonds between the drugs, agents or compounds may be utilized. In
addition, various materials may be utilized as surface modifications to
prevent the drugs, agents or compounds from coming off prematurely.
BRIEF DESCRIPTION OF THE DRAWINGS
[0039] The foregoing and other features and advantages of the invention
will be apparent from the following, more particular description of
preferred embodiments of the invention, as illustrated in the
accompanying drawings.
[0040] FIG. 1 is a view along the length of a stent (ends not shown) prior
to expansion showing the exterior surface of the stent and the
characteristic banding pattern.
[0041] FIG. 2 is a view along the length of the stent of FIG. 1 having
reservoirs in accordance with the present invention.
[0042] FIG. 3 indicates the fraction of drug released as a function of
time from coatings of the present invention over which no topcoat has
been disposed.
[0043] FIG. 4 indicates the fraction of drug released as a function of
time from coatings of the present invention including a topcoat disposed
thereon.
[0044] FIG. 5 indicates the fraction of drug released as a function of
time from coatings of the present invention over which no topcoat has
been disposed.
[0045] FIG. 6 indicates in vivo stent release kinetics of rapamycin from
poly(VDF/HFP).
[0046] FIG. 7 is a cross-sectional view of a band of the stent of FIG. 1
having drug coatings thereon in accordance with a first exemplary
embodiment of the invention.
[0047] FIG. 8 is a cross-sectional view of a band of the stent of FIG. 1
having drug coatings thereon in accordance with a second exemplary
embodiment of the invention.
[0048] FIG. 9 is a cross-sectional view of a band of the stent of FIG. 1
having drug coatings thereon in accordance with a third exemplary
embodiment of the present invention.
[0049] FIG. 10 is a perspective view of an exemplary stent in its
compressed state which may be utilized in conjunction with the present
invention.
[0050] FIG. 11 is a sectional, flat view of the stent shown in FIG. 10.
[0051] FIG. 12 is a perspective view of the stent shown in FIG. 10 but
showing it in its expanded state.
[0052] FIG. 13 is an enlarged sectional view of the stent shown in FIG.
12.
[0053] FIG. 14 is an enlarged view of section of the stent shown in FIG.
11.
[0054] FIG. 15 is a view similar to that of FIG. 11 but showing an
alternate embodiment of the stent.
[0055] FIG. 16 is a perspective view of the stent of FIG. 10 having a
plurality of markers attached to the ends thereof in accordance with the
present invention.
[0056] FIG. 17 is a cross-sectional view of a marker in accordance with
the present invention.
[0057] FIG. 18 is an enlarged perspective view of an end of the stent with
the markers forming a substantially straight line in accordance with the
present invention.
[0058] FIG. 19 is a simplified partial cross-sectional view of a stent
delivery apparatus having a stent loaded therein, which can be used with
a stent made in accordance with the present invention.
[0059] FIG. 20 is a view similar to that of FIG. 19 but showing an
enlarged view of the distal end of the apparatus.
[0060] FIG. 21 is a perspective view of an end of the stent with the
markers in a partially expanded form as it emerges from the delivery
apparatus in accordance with the present invention.
[0061] FIG. 22 is a cross-sectional view of a balloon having a lubricious
coating affixed thereto in accordance with the present invention.
[0062] FIG. 23 is a cross-sectional view of a band of the stent in FIG. 1
having a lubricious coating affixed thereto in accordance with the
present invention.
[0063] FIG. 24 is a cross-sectional view of a self-expanding stent in a
delivery device having a lubricious coating in accordance with the
present invention.
[0064] FIG. 25 is a cross-sectional view of a band of the stent in FIG. 1
having a modified polymer coating in accordance with the present
invention.
[0065] FIG. 26 illustrates an exemplary balloon-expandable stent having an
alternative arrangement of "N" and "J" links between sets of strut
members, represented on a flat, two-dimensional plan view in accordance
with the present invention.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0066] The drug/drug combinations and delivery devices of the present
invention may be utilized to effectively prevent and treat vascular
disease, and in particular, vascular disease caused by injury. Various
medical treatment devices utilized in the treatment of vascular disease
may ultimately induce further complications. For example, balloon
angioplasty is a procedure utilized to increase blood flow through an
artery and is the predominant treatment for coronary vessel stenosis.
However, as stated above, the procedure typically causes a certain degree
of damage to the vessel wall, thereby potentially exacerbating the
problem at a point later in time. Although other procedures and diseases
may cause similar injury, exemplary embodiments of the present invention
will be described with respect to the treatment of restenosis and related
complications following percutaneous transluminal coronary angioplasty
and other similar arterial/venous procedures.
[0067] While exemplary embodiments of the invention will be described with
respect to the treatment of restenosis and related complications
following percutaneous transluminal coronary angioplasty, it is important
to note that the local delivery of drug/drug combinations may be utilized
to treat a wide variety of conditions utilizing any number of medical
devices, or to enhance the function and/or life of the device. For
example, intraocular lenses, placed to restore vision after cataract
surgery is often compromised by the formation of a secondary cataract.
The latter is often a result of cellular overgrowth on the lens surface
and can be potentially minimized by combining a drug or drugs with the
device. Other medical devices which often fail due to tissue in-growth or
accumulation of proteinaceous material in, on and around the device, such
as shunts for hydrocephalus, dialysis grafts, colostomy bag attachment
devices, ear drainage tubes, leads for pace makers and implantable
defibrillators can also benefit from the device-drug combination
approach.
[0068] Devices which serve to improve the structure and function of tissue
or organ may also show benefits when combined with the appropriate agent
or agents. For example, improved osteointegration of orthopedic devices
to enhance stabilization of the implanted device could potentially be
achieved by combining it with agents such as bone-morphogenic protein.
Similarly other surgical devices, sutures, staples, anastomosis devices,
vertebral disks, bone pins, suture anchors, hemostatic barriers, clamps,
screws, plates, clips, vascular implants, tissue adhesives and sealants,
tissue scaffolds, various types of dressings, bone substitutes,
intraluminal devices, and vascular supports could also provide enhanced
patient benefit using this drug-device combination approach. Essentially,
any type of medical device may be coated in some fashion with a drug or
drug combination which enhances treatment over use of the singular use of
the device or pharmaceutical agent.
[0069] In addition to various medical devices, the coatings on these
devices may be used to deliver therapeutic and pharmaceutic agents
including: antiproliferative/antimitotic agents including natural
products such as vinca alkaloids (i.e. vinblastine, vincristine, and
vinorelbine), paclitaxel, epidipodophyllotoxins (i.e. etoposide,
teniposide), antibiotics (dactinomycin (actinomycin D) daunorubicin,
doxorubicin and idarubicin), anthracyclines, mitoxantrone, bleomycins,
plicamycin (mithramycin) and mitomycin, enzymes (L-asparaginase which
systemically metabolizes L-asparagine and deprives cells which do not
have the capacity to synthesize their own asparagine); antiplatelet
agents such as G(GP)II.sub.bIII.sub.a inhibitors and vitronectin receptor
antagonists; antiproliferative/antimitotic alkylating agents such as
nitrogen mustards (mechlorethamine, cyclophosphamide and analogs,
melphalan, chlorambucil), ethylenimines and methylmelamines
(hexamethylmelamine and thiotepa), alkyl sulfonates-busulfan,
nirtosoureas (carmustine (BCNU) and analogs, streptozocin),
trazenes--dacarbazinine (DTIC); antiproliferative/antimitotic
antimetabolites such as folic acid analogs (methotrexate), pyrimidine
analogs (fluorouracil, floxuridine, and cytarabine), purine analogs and
related inhibitors (mercaptopurine, thioguanine, pentostatin and
2-chlorodeoxyadenosine {cladribine}); platinum coordination complexes
(cisplatin, carboplatin), procarbazine, hydroxyurea, mitotane,
aminoglutethimide; hormones (i.e. estrogen); anticoagulants (heparin,
synthetic heparin salts and other inhibitors of thrombin); fibrinolytic
agents (such as tissue plasminogen activator, streptokinase and
urokinase), aspirin, dipyridamole, ticlopidine, clopidogrel, abciximab;
antimigratory; antisecretory (breveldin); antiinflammatory: such as
adrenocortical steroids (cortisol, cortisone, fludrocortisone,
prednisone, prednisolone, 6.alpha.-methylprednisolone, triamcinolone,
betamethasone, and dexamethasone), non-steroidal agents (salicylic acid
derivatives i.e. aspirin; para-aminophenol derivatives i.e.
acetominophen; indole and indene acetic acids (indomethacin, sulindac,
and etodalac), heteroaryl acetic acids (tolmetin, diclofenac, and
ketorolac), arylpropionic acids (ibuprofen and derivatives), anthranilic
acids (mefenamic acid, and meclofenamic acid), enolic acids (piroxicam,
tenoxicam, phenylbutazone, and oxyphenthatrazone), nabumetone, gold
compounds (auranofin, aurothioglucose, gold sodium thiomalate);
immunosuppressives: (cyclosporine, tacrolimus (FK-506), sirolimus
(rapamycin), azathioprine, mycophenolate mofetil); angiogenic agents:
vascular endothelial growth factor (VEGF), fibroblast growth factor
(FGF); angiotensin receptor blocker; nitric oxide donors; anti-sense
oligionucleotides and combinations thereof; cell cycle inhibitors, mTOR
inhibitors, and growth factor signal transduction kinase inhibitors.
[0070] As stated previously, the implantation of a coronary stent in
conjunction with balloon angioplasty is highly effective in treating
acute vessel closure and may reduce the risk of restenosis. Intravascular
ultrasound studies (Mintz et al., 1996) suggest that coronary stenting
effectively prevents vessel constriction and that most of the late
luminal loss after stent implantation is due to plaque growth, probably
related to neointimal hyperplasia. The late luminal loss after coronary
stenting is almost two times higher than that observed after conventional
balloon angioplasty. Thus, inasmuch as stents prevent at least a portion
of the restenosis process, a combination of drugs, agents or compounds
which prevents smooth muscle cell proliferation, reduces inflammation and
reduces coagulation or prevents smooth muscle cell proliferation by
multiple mechanisms, reduces inflammation and reduces coagulation
combined with a stent may provide the most efficacious treatment for
post-angioplasty restenosis. The systemic use of drugs, agents or
compounds in combination with the local delivery of the same or different
drug/drug combinations may also provide a beneficial treatment option.
[0071] The local delivery of drug/drug combinations from a stent has the
following advantages; namely, the prevention of vessel recoil and
remodeling through the scaffolding action of the stent and the prevention
of multiple components of neointimal hyperplasia or restenosis as well as
a reduction in inflammation and thrombosis. This local administration of
drugs, agents or compounds to stented coronary arteries may also have
additional therapeutic benefit. For example, higher tissue concentrations
of the drugs, agents or compounds may be achieved utilizing local
delivery, rather than systemic administration. In addition, reduced
systemic toxicity may be achieved utilizing local delivery rather than
systemic administration while maintaining higher tissue concentrations.
Also in utilizing local delivery from a stent rather than systemic
administration, a single procedure may suffice with better patient
compliance. An additional benefit of combination drug, agent, and/or
compound therapy may be to reduce the dose of each of the therapeutic
drugs, agents or compounds, thereby limiting their toxicity, while still
achieving a reduction in restenosis, inflammation and thrombosis. Local
stent-based therapy is therefore a means of improving the therapeutic
ratio (efficacy/toxicity) of anti-restenosis, anti-inflammatory,
anti-thrombotic drugs, agents or compounds.
[0072] There are a multiplicity of different stents that may be utilized
following percutaneous transluminal coronary angioplasty. Although any
number of stents may be utilized in accordance with the present
invention, for simplicity, a limited number of stents will be described
in exemplary embodiments of the present invention. The skilled artisan
will recognize that any number of stents may be utilized in connection
with the present invention. In addition, as stated above, other medical
devices may be utilized.
[0073] A stent is commonly used as a tubular structure left inside the
lumen of a duct to relieve an obstruction. Commonly, stents are inserted
into the lumen in a non-expanded form and are then expanded autonomously,
or with the aid of a second device in situ. A typical method of expansion
occurs through the use of a catheter-mounted angioplasty balloon which is
inflated within the stenosed vessel or body passageway in order to shear
and disrupt the obstructions associated with the wall components of the
vessel and to obtain an enlarged lumen.
[0074] FIG. 1 illustrates an exemplary stent 100 which may be utilized in
accordance with an exemplary embodiment of the present invention. The
expandable cylindrical stent 100 comprises a fenestrated structure for
placement in a blood vessel, duct or lumen to hold the vessel, duct or
lumen open, more particularly for protecting a segment of artery from
restenosis after angioplasty. The stent 100 may be expanded
circumferentially and maintained in an expanded configuration, that is
circumferentially or radially rigid. The stent 100 is axially flexible
and when flexed at a band, the stent 100 avoids any externally-protruding
component parts.
[0075] The stent 100 generally comprises first and second ends with an
intermediate section therebetween. The stent 100 has a longitudinal axis
and comprises a plurality of longitudinally disposed bands 102, wherein
each band 102 defines a generally continuous wave along a line segment
parallel to the longitudinal axis. A plurality of circumferentially
arranged links 104 maintain the bands 102 in a substantially tubular
structure. Essentially, each longitudinally disposed band 102 is
connected at a plurality of periodic locations, by a short
circumferentially arranged link 104 to an adjacent band 102. The wave
associated with each of the bands 102 has approximately the same
fundamental spatial frequency in the intermediate section, and the bands
102 are so disposed that the wave associated with them are generally
aligned so as to be generally in phase with one another. As illustrated
in the figure, each longitudinally arranged band 102 undulates through
approximately two cycles before there is a link to an adjacent band 102.
[0076] The stent 100 may be fabricated utilizing any number of methods.
For example, the stent 100 may be fabricated from a hollow or formed
stainless steel tube that may be machined using lasers, electric
discharge milling, chemical etching or other means. The stent 100 is
inserted into the body and placed at the desired site in an unexpanded
form. In one exemplary embodiment, expansion may be effected in a blood
vessel by a balloon catheter, where the final diameter of the stent 100
is a function of the diameter of the balloon catheter used.
[0077] It should be appreciated that a stent 100 in accordance with the
present invention may be embodied in a shape-memory material, including,
for example, an appropriate alloy of nickel and titanium or stainless
steel. Structures formed from stainless steel may be made self-expanding
by configuring the stainless steel in a predetermined manner, for
example, by twisting it into a braided configuration. In this embodiment
after the stent 100 has been formed it may be compressed so as to occupy
a space sufficiently small as to permit its insertion in a blood vessel
or other tissue by insertion means, wherein the insertion means include a
suitable catheter, or flexible rod. On emerging from the catheter, the
stent 100 may be configured to expand into the desired configuration
where the expansion is automatic or triggered by a change in pressure,
temperature or electrical stimulation.
[0078] FIG. 2 illustrates an exemplary embodiment of the present invention
utilizing the stent 100 illustrated in FIG. 1. As illustrated, the stent
100 may be modified to comprise one or more reservoirs 106. Each of the
reservoirs 106 may be opened or closed as desired. These reservoirs 106
may be specifically designed to hold the drug/drug combinations to be
delivered. Regardless of the design of the stent 100, it is preferable to
have the drug/drug combination dosage applied with enough specificity and
a sufficient concentration to provide an effective dosage in the lesion
area. In this regard, the reservoir size in the bands 102 is preferably
sized to adequately apply the drug/drug combination dosage at the desired
location and in the desired amount.
[0079] In an alternate exemplary embodiment, the entire inner and outer
surface of the stent 100 may be coated with drug/drug combinations in
therapeutic dosage amounts. A detailed description of a drug for treating
restenosis, as well as exemplary coating techniques, is described below.
It is, however, important to note that the coating techniques may vary
depending on the drug/drug combinations. Also, the coating techniques may
vary depending on the material comprising the stent or other intraluminal
medical device.
[0080] FIG. 26 illustrates another exemplary embodiment of a
balloon-expandable stent. FIG. 26 illustrates the stent 900 in its
crimped, pre-deployed state as it would appear if it were cut
longitudinally and then laid out into a flat, two-dimensional
configuration. The stent 900 has curved end struts 902 and diagonal
struts 904 with each set of strut members 906 connected by sets of
flexible links 908, 910 or 912. In this exemplary embodiment, three
different types of flexible links are used. A set of "N" links 910
comprising six circumferentially spaced "N" links 914 and a set of
inverted "N" links 912 comprising six circumferentially spaced inverted
"N" links 916 each connect to adjacent sets of strut members 906 at the
ends of the stent 900. A set of inverted "J" links 918 comprising six
circumferentially spaced inverted "J" links 908 are used to connect the
adjacent sets of strut members 906 in the center of the stent 900. The
shape of the "N" links 914 and inverted "N" links 916 facilitate the
links' ability to lengthen and shorten as the stent bends around a curve
during delivery into the human body. This ability to lengthen and shorten
helps to prevent the sets of strut members from being pushed or pulled
off the balloon during delivery into the body and is particularly
applicable to short stents which tend to have relatively poor stent
retention onto an inflatable balloon. The stent 900 with its greater
strength at its central region would advantageously be used for
comparatively short stenoses that have a tough, calcified central
section. It should also be understood that a regular "J" link could be
used for the stent 900 in place of the inverted "J" link 908. Other
exemplary embodiments of balloon expandable stents may be found in U.S.
Pat. No. 6,190,403 B1 issued on Feb. 20, 2001 and which is incorporated
by reference herein.
[0081] Rapamycin is a macrocyclic triene antibiotic produced by
Streptomyces hygroscopicus as disclosed in U.S. Pat. No. 3,929,992. It
has been found that rapamycin among other things inhibits the
proliferation of vascular smooth muscle cells in vivo. Accordingly,
rapamycin may be utilized in treating intimal smooth muscle cell
hyperplasia, restenosis, and vascular occlusion in a mammal, particularly
following either biologically or mechanically mediated vascular injury,
or under conditions that would predispose a mammal to suffering such a
vascular injury. Rapamycin functions to inhibit smooth muscle cell
proliferation and does not interfere with the re-endothelialization of
the vessel walls.
[0082] Rapamycin reduces vascular hyperplasia by antagonizing smooth
muscle proliferation in response to mitogenic signals that are released
during an angioplasty induced injury. Inhibition of growth factor and
cytokine mediated smooth muscle proliferation at the late G1 phase of the
cell cycle is believed to be the dominant mechanism of action of
rapamycin. However, rapamycin is also known to prevent T-cell
proliferation and differentiation when administered systemically. This is
the basis for its immunosuppresive activity and its ability to prevent
graft rejection.
[0083] As used herein, rapamycin includes rapamycin and all analogs,
derivatives and congeners that find FKBP12, and other immunophilins, and
possesses the same pharmacologic properties as rapamycin.
[0084] Although the anti-proliferative effects of rapamycin may be
achieved through systemic use, superior results may be achieved through
the local delivery of the compound. Essentially, rapamycin works in the
tissues, which are in proximity to the compound, and has diminished
effect as the distance from the delivery device increases. In order to
take advantage of this effect, one would want the rapamycin in direct
contact with the lumen walls. Accordingly, in a preferred embodiment, the
rapamycin is incorporated onto the surface of the stent or portions
thereof. Essentially, the rapamycin is preferably incorporated into the
stent 100, illustrated in FIG. 1, where the stent 100 makes contact with
the lumen wall.
[0085] Rapamycin may be incorporated onto or affixed to the stent in a
number of ways. In the exemplary embodiment, the rapamycin is directly
incorporated into a polymeric matrix and sprayed onto the outer surface
of the stent. The rapamycin elutes from the polymeric matrix over time
and enters the surrounding tissue. The rapamycin preferably remains on
the stent for at least three days up to approximately six months, and
more preferably between seven and thirty days.
[0086] Any number of non-erodible polymers may be utilized in conjunction
with the rapamycin. In one exemplary embodiment, the polymeric matrix
comprises two layers. The base layer comprises a solution of
poly(ethylene-covinylacetate) and polybutylmethacrylate. The rapamycin is
incorporated into this base layer. The outer layer comprises only
polybutylmethacrylate and acts as a diffusion barrier to prevent the
rapamycin from eluting too quickly. The thickness of the outer layer or
top coat determines the rate at which the rapamycin elutes from the
matrix. Essentially, the rapamycin elutes from the matrix by diffusion
through the polymer matrix. Polymers are permeable, thereby allowing
solids, liquids and gases to escape therefrom. The total thickness of the
polymeric matrix is in the range from about one micron to about twenty
microns or greater. It is important to note that primer layers and metal
surface treatments may be utilized before the polymeric matrix is affixed
to the medical device. For example, acid cleaning, alkaline (base)
cleaning, salinization and parylene deposition may be used as part of the
overall process described below.
[0087] The poly(ethylene-co-vinylacetate), polybutylmethacrylate and
rapamycin solution may be incorporated into or onto the stent in a number
of ways. For example, the solution may be sprayed onto the stent or the
stent may be dipped into the solution. Other methods include spin coating
and RF-plasma polymerization. In one exemplary embodiment, the solution
is sprayed onto the stent and then allowed to dry. In another exemplary
embodiment, the solution may be electrically charged to one polarity and
the stent electrically changed to the opposite polarity. In this manner,
the solution and stent will be attracted to one another. In using this
type of spraying process, waste may be reduced and more precise control
over the thickness of the coat may be achieved.
[0088] In another exemplary embodiment, the rapamycin or other therapeutic
agent may be incorporated into a film-forming polyfluoro copolymer
comprising an amount of a first moiety selected from the group consisting
of polymerized vinylidenefluoride and polymerized tetrafluoroethylene,
and an amount of a second moiety other than the first moiety and which is
copolymerized with the first moiety, thereby producing the polyfluoro
copolymer, the second moiety being capable of providing toughness or
elastomeric properties to the polyfluoro copolymer, wherein the relative
amounts of the first moiety and the second moiety are effective to
provide the coating and film produced therefrom with properties effective
for use in treating implantable medical devices.
[0089] The present invention provides polymeric coatings comprising a
polyfluoro copolymer and implantable medical devices, for example, stents
coated with a film of the polymeric coating in amounts effective to
reduce thrombosis and/or restenosis when such stents are used in, for
example, angioplasty procedures. As used herein, polyfluoro copolymers
means those copolymers comprising an amount of a first moiety selected
from the group consisting of polymerized vinylidenefluoride and
polymerized tetrafluoroethylene, and an amount of a second moiety other
than the first moiety and which is copolymerized with the first moiety to
produce the polyfluoro copolymer, the second moiety being capable of
providing toughness or elastomeric properties to the polyfluoro
copolymer, wherein the relative amounts of the first moiety and the
second moiety are effective to provide coatings and film made from such
polyfluoro copolymers with properties effective for use in coating
implantable medical devices.
[0090] The coatings may comprise pharmaceutical or therapeutic agents for
reducing restenosis, inflammation and/or thrombosis, and stents coated
with such coatings may provide sustained release of the agents. Films
prepared from certain polyfluoro copolymer coatings of the present
invention provide the physical and mechanical properties required of
conventional coated medical devices, even where maximum temperature, to
which the device coatings and films are exposed, are limited to
relatively low temperatures. This is particularly important when using
the coating/film to deliver pharmaceutical/therapeutic agents or drugs
that are heat sensitive, or when applying the coating onto
temperature-sensitive devices such as catheters. When maximum exposure
temperature is not an issue, for example, where heat-stable agents such
as itraconazole are incorporated into the coatings, higher melting
thermoplastic polyfluoro copolymers may be used and, if very high
elongation and adhesion is required, elastomers may be used. If desired
or required, the polyfluoro elastomers may be crosslinked by standard
methods described in, e.g., Modern Fluoropolymers, (J. Shires ed.) John
Wiley & Sons, New York, 1997, pp. 77-87.
[0091] The present invention comprises polyfluoro copolymers that provide
improved biocompatible coatings or vehicles for medical devices. These
coatings provide inert biocompatible surfaces to be in contact with body
tissue of a mammal, for example, a human, sufficient to reduce
restenosis, or thrombosis, or other undesirable reactions. While many
reported coatings made from polyfluoro homopolymers are insoluble and/or
require high heat, for example, greater than about one hundred
twenty-five degrees centigrade, to obtain films with adequate physical
and mechanical properties for use on implantable devices, for example,
stents, or are not particularly tough or elastomeric, films prepared from
the polyfluoro copolymers of the present invention provide adequate
adhesion, toughness or elasticity, and resistance to cracking when formed
on medical devices. In certain exemplary embodiments, this is the case
even where the devices are subjected to relatively low maximum
temperatures.
[0092] The polyfluoro copolymers used for coatings according to the
present invention are preferably film-forming polymers that have
molecular weight high enough so as not to be waxy or tacky. The polymers
and films formed therefrom should preferably adhere to the stent and not
be readily deformable after deposition on the stent as to be able to be
displaced by hemodynamic stresses. The polymer molecular weight should
preferably be high enough to provide sufficient toughness so that films
comprising the polymers will not be rubbed off during handling or
deployment of the stent. In certain exemplary embodiments the coating
will not crack where expansion of the stent or other medical devices
occurs.
[0093] Coatings of the present invention comprise polyfluoro copolymers,
as defined hereinabove. The second moiety polymerized with the first
moiety to prepare the polyfluoro copolymer may be selected from those
polymerized, biocompatible monomers that would provide biocompatible
polymers acceptable for implantation in a mammal, while maintaining
sufficient elastomeric film properties for use on medical devices claimed
herein. Such monomers include, without limitation, hexafluoropropylene
(HFP), tetrafluoroethylene (TFE), vinylidenefluoride,
1-hydropentafluoropropylene, perfluoro(methyl vinyl ether),
chlorotrifluoroethylene (CTFE), pentafluoropropene, trifluoroethylene,
hexafluoroacetone and hexafluoroisobutylene.
[0094] Polyfluoro copolymers used in the present invention typically
comprise vinylidinefluoride copolymerized with hexafluoropropylene, in
the weight ratio in the range of from about fifty to about ninety-two
weight percent vinylidinefluoride to about fifty to about eight weight
percent HFP. Preferably, polyfluoro copolymers used in the present
invention comprise from about fifty to about eighty-five weight percent
vinylidinefluoride copolymerized with from about fifty to about fifteen
weight percent HFP. More preferably, the polyfluoro copolymers will
comprise from about fifty-five to about seventy weight percent
vinylidineflyoride copolymerized with from about forty-five to about
thirty weight percent HFP. Even more preferably, polyfluoro copolymers
comprise from about fifty-five to about sixty-five weight percent
vinylidinefluoride copolymerized with from about forty-five to about
thirty-five weight percent HFP. Such polyfluoro copolymers are soluble,
in varying degrees, in solvents such as dimethylacetamide (DMAc),
tetrahydrofuran, dimethyl formamide, dimethyl sulfoxide and n-methyl
pyrrolidone. Some are soluble in methylethylketone (MEK), acetone,
methanol and other solvents commonly used in applying coatings to
conventional implantable medical devices.
[0095] Conventional polyfluoro homopolymers are crystalline and difficult
to apply as high quality films onto metal surfaces without exposing the
coatings to relatively high temperatures that correspond to the melting
temperature (Tm) of the polymer. The elevated temperature serves to
provide films prepared from such PVDF homopolymer coatings that exhibit
sufficient adhesion of the film to the device, while preferably
maintaining sufficient flexibility to resist film cracking upon
expansion/contraction of the coated medical device. Certain films and
coatings according to the present invention provide these same physical
and mechanical properties, or essentially the same properties, even when
the maximum temperatures to which the coatings and films are exposed is
less than about a maximum predetermined temperature. This is particularly
important when the coatings/films comprise pharmaceutical or therapeutic
agents or drugs that are heat sensitive, for example, subject to chemical
or physical degradation or other heat-induced negative affects, or when
coating heat sensitive substrates of medical devices, for example,
subject to heat-induced compositional or structural degradation.
[0096] Depending on the particular device upon which the coatings and
films of the present invention are to be applied and the particular
use/result required of the device, polyfluoro copolymers used to prepare
such devices may be crystalline, semi-crystalline or amorphous.
[0097] Where devices have no restrictions or limitations with respect to
exposure of same to elevated temperatures, crystalline polyfluoro
copolymers may be employed. Crystalline polyfluoro copolymers tend to
resist the tendency to flow under applied stress or gravity when exposed
to temperatures above their glass transition (Tg) temperatures.
Crystalline polyfluoro copolymers provide tougher coatings and films than
their fully amorphous counterparts. In addition, crystalline polymers are
more lubricious and more easily handled through crimping and transfer
processes used to mount self-expanding stents, for example, nitinol
stents.
[0098] Semi-crystalline and amorphous polyfluoro copolymers are
advantageous where exposure to elevated temperatures is an issue, for
example, where heat-sensitive pharmaceutical or therapeutic agents are
incorporated into the coatings and films, or where device design,
structure and/or use preclude exposure to such elevated temperatures.
Semi-crystalline polyfluoro copolymer elastomers comprising relatively
high levels, for example, from about thirty to about forty-five weight
percent of the second moiety, for example, HFP, copolymerized with the
first moiety, for example, VDF, have the advantage of reduced coefficient
of friction and self-blocking relative to amorphous polyfluoro copolymer
elastomers. Such characteristics may be of significant value when
processing, packaging and delivering medical devices coated with such
polyfluoro copolymers. In addition, such polyfluoro copolymer elastomers
comprising such relatively high content of the second moiety serves to
control the solubility of certain agents, for example, rapamycin, in the
polymer and therefore controls permeability of the agent through the
matrix.
[0099] Polyfluoro copolymers utilized in the present inventions may be
prepared by various known polymerization methods. For example, high
pressure, free-radical, semi-continuous emulsion polymerization
techniques such as those disclosed in Fluoroelastomers-dependence of
relaxation phenomena on compositions, POLYMER 30, 2180, 1989, by Ajroldi,
et al., may be employed to prepare amorphous polyfluoro copolymers, some
of which may be elastomers. In addition, free-radical batch emulsion
polymerization techniques disclosed herein may be used to obtain polymers
that are semi-crystalline, even where relatively high levels of the
second moiety are included.
[0100] As described above, stents may comprise a wide variety of materials
and a wide variety of geometries. Stents may be made of biocomptible
materials, including biostable and bioabsorbable materials. Suitable
biocompatible metals include, but are not limited to, stainless steel,
tantalum, titanium alloys (including nitinol), and cobalt alloys
(including cobalt-chromium nickel alloys). Suitable nonmetallic
biocompatible materials include, but are not limited to, polyamides,
polyolefins (i.e. polypropylene, polyethylene etc.), nonabsorbable
polyesters (i.e. polyethylene terephthalate), and bioabsorbable aliphatic
polyesters (i.e. homopolymers and copolymers of lactic acid, glycolic
acid, lactide, glycolide, para-dioxanone, trimethylene carbonate,
.epsilon.-caprolactone, and blends thereof).
[0101] The film-forming biocompatible polymer coatings generally are
applied to the stent in order to reduce local turbulence in blood flow
through the stent, as well as adverse tissue reactions. The coatings and
films formed therefrom also may be used to administer a pharmaceutically
active material to the site of the stent placement. Generally, the amount
of polymer coating to be applied to the stent will vary depending on,
among other possible parameters, the particular polyfluoro copolymer used
to prepare the coating, the stent design and the desired effect of the
coating. Generally, the coated stent will comprise from about 0.1 to
about fifteen weight percent of the coating, preferably from about 0.4 to
about ten weight percent. The polyfluoro copolymer coatings may be
applied in one or more coating steps, depending on the amount of
polyfluoro copolymer to be applied. Different polyfluoro copolymers may
be used for different layers in the stent coating. In fact, in certain
exemplary embodiments, it is highly advantageous to use a diluted first
coating solution comprising a polyfluoro copolymer as a primer to promote
adhesion of a subsequent polyfluoro copolymer coating layer that may
include pharmaceutically active materials. The individual coatings may be
prepared from different polyfluoro copolymers.
[0102] Additionally, a top coating may be applied to delay release of the
pharmaceutical agent, or they could be used as the matrix for the
delivery of a different pharmaceutically active material. Layering of
coatings may be used to stage release of the drug or to control release
of different agents placed in different layers.
[0103] Blends of polyfluoro copolymers may also be used to control the
release rate of different agents or to provide a desirable balance of
coating properties, i.e. elasticity, toughness, etc., and drug delivery
characteristics, for example, release profile. Polyfluoro copolymers with
different solubilities in solvents may be used to build up different
polymer layers that may be used to deliver different drugs or to control
the release profile of a drug. For example, polyfluoro copolymers
comprising 85.5/14.5 (wt/wt) of poly(vinylidinefluoride/HFP) and
60.6/39.4 (wt/wt) of poly(vinylidinefluoride/HFP) are both soluble in
DMAc. However, only the 60.6/39.4 PVDF polyfluoro copolymer is soluble in
methanol. So, a first layer of the 85.5/14.5 PVDF polyfluoro copolymer
comprising a drug could be over coated with a topcoat of the 60.6/39.4
PVDF polyfluoro copolymer made with the methanol solvent. The top coating
may be used to delay the drug delivery of the drug contained in the first
layer. Alternately, the second layer could comprise a different drug to
provide for sequential drug delivery. Multiple layers of different drugs
could be provided by alternating layers of first one polyfluoro
copolymer, then the other. As will be readily appreciated by those
skilled in the art, numerous layering approaches may be used to provide
the desired drug delivery.
[0104] Coatings may be formulated by mixing one or more therapeutic agents
with the coating polyfluoro copolymers in a coating mixture. The
therapeutic agent may be present as a liquid, a finely divided solid, or
any other appropriate physical form. Optionally, the coating mixture may
include one or more additives, for example, nontoxic auxiliary substances
such as diluents, carriers, excipients, stabilizers or the like. Other
suitable additives may be formulated with the polymer and
pharmaceutically active agent or compound. For example, a hydrophilic
polymer may be added to a biocompatible hydrophobic coating to modify the
release profile, or a hydrophobic polymer may be added to a hydrophilic
coating to modify the release profile. One example would be adding a
hydrophilic polymer selected from the group consisting of polyethylene
oxide, polyvinyl pyrrolidone, polyethylene glycol, carboxylmethyl
cellulose, and hydroxymethyl cellulose to a polyfluoro copolymer coating
to modify the release profile. Appropriate relative amounts may be
determined by monitoring the in vitro and/or in vivo release profiles for
the therapeutic agents.
[0105] The best conditions for the coating application are when the
polyfluoro copolymer and pharmaceutic agent have a common solvent. This
provides a wet coating that is a true solution. Less desirable, yet still
usable, are coatings that contain the pharmaceutical agent as a solid
dispersion in a solution of the polymer in solvent. Under the dispersion
conditions, care must be taken to ensure that the particle size of the
dispersed pharmaceutical powder, both the primary powder size and its
aggregates and agglomerates, is small enough not to cause an irregular
coating surface or to clog the slots of the stent that need to remain
essentially free of coating. In cases where a dispersion is applied to
the stent and the smoothness of the coating film surface requires
improvement, or to be ensured that all particles of the drug are fully
encapsulated in the polymer, or in cases where the release rate of the
drug is to be slowed, a clear (polyfluoro copolymer only) topcoat of the
same polyfluoro copolymer used to provide sustained release of the drug
or another polyfluoro copolymer that further restricts the diffusion of
the drug out of the coating may be applied. The topcoat may be applied by
dip coating with mandrel to clear the slots. This method is disclosed in
U.S. Pat. No. 6,153,252. Other methods for applying the topcoat include
spin coating and spray coating. Dip coating of the topcoat can be
problematic if the drug is very soluble in the coating solvent, which
swells the polyfluoro copolymer, and the clear coating solution acts as a
zero concentration sink and redissolves previously deposited drug. The
time spent in the dip bath may need to be limited so that the drug is not
extracted out into the drug-free bath. Drying should be rapid so that the
previously deposited drug does not completely diffuse into the topcoat.
[0106] The amount of therapeutic agent will be dependent upon the
particular drug employed and medical condition being treated. Typically,
the amount of drug represents about 0.001 percent to about seventy
percent, more typically about 0.001 percent to about sixty percent.
[0107] The quantity and type of polyfluoro copolymers employed in the
coating film comprising the pharmaceutic agent will vary depending on the
release profile desired and the amount of drug employed. The product may
contain blends of the same or different polyfluoro copolymers having
different molecular weights to provide the desired release profile or
consistency to a given formulation.
[0108] Polyfluoro copolymers may release dispersed drug by diffusion. This
can result in prolonged delivery (over, say approximately one to
two-thousand hours, preferably two to eight-hundred hours) of effective
amounts (0.001 .mu.g/cm.sup.2-min to 1000 .mu.g/cm.sup.2-min) of the
drug. The dosage may be tailored to the subject being treated, the
severity of the affliction, the judgment of the prescribing physician,
and the like.
[0109] Individual formulations of drugs and polyfluoro copolymers may be
tested in appropriate in vitro and in vivo models to achieve the desired
drug release profiles. For example, a drug could be formulated with a
polyfluoro copolymer, or blend of polyfluoro copolymers, coated onto a
stent and placed in an agitated or circulating fluid system, for example,
twenty-five percent ethanol in water. Samples of the circulating fluid
could be taken to determine the release profile (such as by HPLC, UV
analysis or use of radiotagged molecules). The release of a
pharmaceutical compound from a stent coating into the interior wall of a
lumen could be modeled in appropriate animal system. The drug release
profile could then be monitored by appropriate means such as, by taking
samples at specific times and assaying the samples for drug concentration
(using HPLC to detect drug concentration). Thrombus formation can be
modeled in animal models using the In-platelet imaging methods described
by Hanson and Harker, Proc. Natl. Acad. Sci. USA 85:3184-3188 (1988).
Following this or similar procedures, those skilled in the art will be
able to formulate a variety of stent coating formulations.
[0110] While not a requirement of the present invention, the coatings and
films may bp crosslinked once applied to the medical devices.
Crosslinking may be affected by any of the known crosslinking mechanisms,
such as chemical, heat or light. In addition, crosslinking initiators and
promoters may be used where applicable and appropriate. In those
exemplary embodiments utilizing crosslinked films comprising
pharmaceutical agents, curing may affect the rate at which the drug
diffuses from the coating. Crosslinked polyfluoro copolymers films and
coatings of the present invention also may be used without drug to modify
the surface of implantable medical devices. 25
EXAMPLES
Example 1
[0111] A PVDF homopolymer (Solef.RTM. 1008 from Solvay Advanced Polymers,
Houston, Tex., Tm about 175.degree. C.) and polyfluoro copolymers of
poly(vinylidenefluoride/HFP), 92/8 and 91/9 weight percent
vinylidenefluoride/HFP as determined by F.sup.9 NMR, respectively (eg:
Solef.RTM. 11010 and 11008, Solvay Advanced Polymers, Houston, Tex., Tm
about 159 degrees C. and 160 degrees C., respectively) were examined as
potential coatings for stents. These polymers are soluble in solvents
such as, but not limited to, DMAc, N,N-dimethylformamide (DMF), dimethyl
sulfoxide (DMSO), N-methylpyrrolidone (NMP), tetrahydrofuran (THF) and
acetone. Polymer coatings were prepared by dissolving the polymers in
acetone, at five weight percent as a primer, or by dissolving the polymer
in 50/50 DMAc/acetone, at thirty weight percent as a topcoat. Coatings
that were applied to the stents by dipping and dried at 60 degrees C. in
air for several hours, followed by 60 degrees C. for three hours in a
<100 mm Hg vacuum, resulted in white foamy films. As applied, these
films adhered poorly to the stent and flaked off, indicating they were
too brittle. When stents coated in this manner were heated above 175
degrees C., i.e. above the melting temperature of the polymer, a clear,
adherent film was formed. Since coatings require high temperatures, for
example, above the melting temperature of the polymer, to achieve high
quality films. As mentioned above, the high temperature heat treatment is
unacceptable for the majority of drug compounds due to their thermal
sensitivity.
Example 2
[0112] A polyfluoro copolymer (Solef.RTM. 21508) comprising 85.5 weight
percent vinylidenefluoride copolymerized with 14.5 weight percent HFP, as
determined by F.sup.19 NMR, was evaluated. This copolymer is less
crystalline than the polyfluoro homopolymer and copolymers described in
Example 1. It also has a lower melting point reported to be about 133
degrees C. Once again, a coating comprising about twenty weight percent
of the polyfluoro copolymer was applied from a polymer solution in 50/50
DMAc/MEK. After drying (in air) at 60 degrees C. for several hours,
followed by 60 degrees C. for three hours in a <100 mtorr Hg vacuum,
clear adherent films were obtained. This eliminated the need for a high
temperature heat treatment to achieve high quality films. Coatings were
smoother and more adherent than those of Example 1. Some coated stents
that underwent expansion show some degree of adhesion loss and "tenting"
as the film pulls away from the metal. Where necessary, modification of
coatings containing such copolymers may be made, e.g. by addition of
plasticizers or the like to the coating compositions. Films prepared from
such coatings may be used to coat stents or other medical devices,
particularly where those devices are not susceptible to expansion to the
degree of the stents.
[0113] The coating process above was repeated, this time with a coating
comprising the 85.5/14.6 (wt/wt) (vinylidenefluoride/HFP) and about
thirty weight percent of rapamycin (Wyeth-Ayerst Laboratories,
Philadelphia, Pa.), based on total weight of coating solids. Clear films
that would occasionally crack or peel upon expansion of the coated stents
resulted. It is believed that inclusion of plasticizers and the like in
the coating composition will result in coatings and films for use on
stents and other medical devices that are not susceptible to such
cracking and peeling.
Example 3
[0114] Polyfluoro copolymers of still higher HFP content were then
examined. This series of polymers were not semicrystalline, but rather
are marketed as elastomers. One such copolymer is Fluorel.TM. FC2261Q
(from Dyneon, a 3M-Hoechst Enterprise, Oakdale, Minn.), a 60.6/39.4
(wt/wt) copolymer of vinylidenefluoride/HFP. Although this copolymer has
a Tg well below room temperature (Tg about minus twenty degrees C.) it is
not tacky at room temperature or even at sixty degrees C. This polymer
has no detectable crystallinity when measured by Differential Scanning
Calorimetry (DSC) or by wide angle X-ray diffraction. Films formed on
stents as described above were non-tacky, clear, and expanded without
incident when the stents were expanded.
[0115] The coating process above was repeated, this time with coatings
comprising the 60.6/39.4 (wt/wt) (vinylidenefluoride/HFP) and about nine,
thirty and fifty weight percent of rapamycin (Wyeth-Ayerst Laboratories,
Philadelphia, Pa.), based on total weight of coating solids,
respectively. Coatings comprising about nine and thirty weight percent
rapamycin provided white, adherent, tough films that expanded without
incident on the stent. Inclusion of fifty percent drug, in the same
manner, resulted in some loss of adhesion upon expansion.
[0116] Changes in the comonomer composition of the polyfluoro copolymer
also can affect the nature of the solid state coating, once dried. For
example, the semicrystalline copolymer, Solef.RTM. 21508, containing 85.5
percent vinylidenefluoride polymerized with 14.5 percent by weight HFP
forms homogeneous solutions with about 30 percent rapamycin (drug weight
divided by total solids weight, for example, drug plus copolymer) in DMAc
and 50/50 DMAc/MEK. When the film is dried (60 degrees C./16 hours
followed by 60 degrees C./3 hours in vacuum of 100 mm Hg) a clear
coating, indicating a solid solution of the drug in the polymer, is
obtained. Conversely, when an amorphous copolymer, Fluorel.TM. FC2261Q,
of PDVF/HFP at 60.6/39.5 (wt/wt) forms a similar thirty percent solution
of rapamycin in DMAc/MEK and is similarly dried, a white film, indicating
phase separation of the drug and the polymer, is obtained. This second
drug containing film is much slower to release the drug into an in vitro
test solution of twenty-five percent ethanol in water than is the former
clear film of crystalline Solef.RTM. 21508. X-ray analysis of both films
indicates that the drug is present in a non-crystalline form. Poor or
very low solubility of the drug in the high HFP containing copolymer
results in slow permeation of the drug through the thin coating film.
Permeability is the product of diffusion rate of the diffusing species
(in this case the drug) through the film (the copolymer) and the
solubility of the drug in the film.
Example 4
In vitro Release Results of Rapamycin from Coating
[0117] FIG. 3 is a plot of data for the 85.5/14.5 vinylidenefluoride/HFP
polyfluoro copolymer, indicating fraction of drug released as a function
of time, with no topcoat. FIG. 4 is a plot of data for the same
polyfluoro copolymer over which a topcoat has been disposed, indicating
that most effect on release rate is with a clear topcoat. As shown
therein, TC150 refers to a device comprising one hundred fifty micrograms
of topcoat, TC235 refers to two hundred thirty-five micrograms of
topcoat, etc. The stents before topcoating had an average of seven
hundred fifty micrograms of coating containing thirty percent rapamycin.
FIG. 5 is a plot for the 60.6/39.4 vinylidenefluoride/HFP polyfluoro
copolymer, indicating fraction of drug released as a function of time,
showing significant control of release rate from the coating without the
use of a topcoat. Release is controlled by loading of drug in the film.
Example 5
In vivo Stent Release Kinetics of Rapamycin from Poly(VDF/HFP)
[0118] Nine New Zealand white rabbits (2.5-3.0 kg) on a normal diet were
given aspirin twenty-four hours prior to surgery, again just prior to
surgery and for the remainder of the study. At the time of surgery,
animals were premedicated with Acepromazine (0.1-0.2 mg/kg) and
anesthetized with a Ketamine/Xylazine mixture (40 mg/kg and 5 mg/kg,
respectively). Animals were given a single intraprocedural dose of
heparin (150 IU/kg, i.v.).
[0119] Arteriectomy of the right common carotid artery was performed and a
5 F catheter introducer (Cordis, Inc.) placed in the vessel and anchored
with ligatures. Iodine contrast agent was injected to visualize the right
common carotid artery, brachlocephalic trunk and aortic arch. A steerable
guide wire (0.014 inch/180 cm, Cordis, Inc.) was inserted via the
introducer and advanced sequentially into each iliac artery to a location
where the artery possesses a diameter closest to 2 mm using the
angiographic mapping done previously. Two stents coated with a film made
of poly(VDF/HFP): (60.6/39.4) with thirty percent rapamycin were deployed
in each animal where feasible, one in each iliac artery, using 3.0 mm
balloon and inflation to 8-10 ATM for thirty seconds followed after a one
minute interval by a second inflation to 8-10 ATM for thirty seconds.
Follow-up angiographs visualizing both iliac arteries are obtained to
confirm correct deployment position of the stent.
[0120] At the end of procedure, the carotid artery was ligated and the
skin is closed with 3/0 vicryl suture using a one layered interrupted
closure. Animals were given butoropanol (0.4 mg/kg, s.c.) and gentamycin
(4 mg/kg, i.m.). Following recovery, the animals were returned to their
cages and allowed free access to food and water.
[0121] Due to early deaths and surgical difficulties, two animals were not
used in this analysis. Stented vessels were removed from the remaining
seven animals at the following time points: one vessel (one animal) at
ten minutes post implant; six vessels (three animals) between forty
minutes and two hours post-implant (average, 1.2 hours); two vessels (two
animals) at three days post implant; and two vessels (one animal) at
seven days post-implant. In one animal at two hours, the stent was
retrieved from the aorta rather than the iliac artery. Upon removal,
arteries were carefully trimmed at both the proximal and distal ends of
the stent. Vessels were then carefully dissected free of the stent,
flushed to remove any residual blood, and both stent and vessel frozen
immediately, wrapped separately in foil, labeled and kept frozen at minus
eighty degrees C. When all samples had been collected, vessels and stents
were frozen, transported and subsequently analyzed for rapamycin in
tissue and results are illustrated in FIG. 4.
Example 6
Purifying the Polymer
[0122] The Fluorel.TM. FC2261Q copolymer was dissolved in MEK at about ten
weight percent and was washed in a 50/50 mixture of ethanol/water at a
14:1 of ethanol/water to the MEK solution ratio. The polymer precipitated
out and was separated from the solvent phase by centrifugation. The
polymer again was dissolved in MEK and the washing procedure repeated.
The polymer was dried after each washing step at sixty degrees C. in a
vacuum oven (<200 mtorr) over night.
[0123] Example 7
In vivo Testing of Coated Stents in Porcine Coronary Arteries
[0124] CrossFlex.RTM. stents (available from Cordis, a Johnson & Johnson
Company) were coated with the "as received" Fluorel.TM. FC2261Q PVDF
copolymer and with the purified polyfluoro copolymer of Example 6, using
the dip and wipe approach. The coated stents were sterilized using
ethylene oxide and a standard cycle. The coated stents and bare metal
stents (controls) were implanted in porcine coronary arteries, where they
remained for twenty-eight days.
[0125] Angiography was performed on the pigs at implantation and at
twenty-eight days. Angiography indicated that the control uncoated stent
exhibited about twenty-one percent restenosis. The polyfluoro copolymer
"as received" exhibited about twenty-six percent restenosis(equivalent to
the control) and the washed copolymer exhibited about 12.5 percent
restenosis.
[0126] Histology results reported neointimal area at twenty-eight days to
be 2.89.+-.0.2, 3.57.+-.0.4 and 2.75.+-.10.3, respectively, for the bare
metal control, the unpurified copolymer and the purified copolymer.
[0127] Since rapamycin acts by entering the surrounding tissue, it is
preferably only affixed to the surface of the stent making contact with
one tissue. Typically, only the outer surface of the stent makes contact
with the tissue. Accordingly, in one exemplary embodiment, only the outer
surface of the stent is coated with rapamycin.
[0128] The circulatory system, under normal conditions, has to be
self-sealing, otherwise continued blood loss from an injury would be life
threatening. Typically, all but the most catastrophic bleeding is rapidly
stopped though a process known as hemostasis. Hemostasis occurs through a
progression of steps. At high rates of flow, hemostasis is a combination
of events involving platelet aggregation and fibrin formation. Platelet
aggregation leads to a reduction in the blood flow due to the formation
of a cellular plug while a cascade of biochemical steps leads to the
formation of a fibrin clot.
[0129] Fibrin clots, as stated above, form in response to injury. There
are certain circumstances where blood clotting or clotting in a specific
area may pose a health risk. For example, during percutaneous
transluminal coronary angioplasty, the endothelial cells of the arterial
walls are typically injured, thereby exposing the sub-endothelial cells.
Platelets adhere to these exposed cells. The aggregating platelets and
the damaged tissue initiate further biochemical process resulting in
blood coagulation. Platelet and fibrin blood clots may prevent the normal
flow of blood to critical areas. Accordingly, there is a need to control
blood clotting in various medical procedures. Compounds that do not allow
blood to clot are called anti-coagulants. Essentially, an anti-coagulant
is an inhibitor of thrombin formation or function. These compounds
include drugs such as heparin and hirudin. As used herein, heparin
includes all direct or indirect inhibitors of thrombin or Factor Xa.
[0130] In addition to being an effective anti-coagulant, heparin has also
been demonstrated to inhibit smooth muscle cell growth in vivo. Thus,
heparin may be effectively utilized in conjunction with rapamycin in the
treatment of vascular disease. Essentially, the combination of rapamycin
and heparin may inhibit smooth muscle cell growth via two different
mechanisms in addition to the heparin acting as an anti-coagulant.
[0131] Because of its multifunctional chemistry, heparin may be
immobilized or affixed to a stent in a number of ways. For example,
heparin may be immobilized onto a variety of surfaces by various methods,
including the photolink methods set forth in U.S. Pat. Nos. 3,959,078 and
4,722,906 to Guire et al. and U.S. Pat. Nos. 5,229,172; 5,308,641;
5,350,800 and 5,415,938 to Cahalan et al. Heparinized surfaces have also
been achieved by controlled release from a polymer matrix, for example,
silicone rubber, as set forth in U.S. Pat. Nos. 5,837,313; 6,099,562 and
6,120,536 to Ding et al.
[0132] In one exemplary embodiment, heparin may be immobilized onto the
stent as briefly described below. The surface onto which the heparin is
to be affixed is cleaned with ammonium peroxidisulfate. Once cleaned,
alternating layers of polyethylenimine and dextran sulfate are deposited
thereon. Preferably, four layers of the polyethylenimine and dextran
sulfate are deposited with a final layer of polyethylenimine.
Aldehyde-end terminated heparin is then immobilized to this final layer
and stabilized with sodium cyanoborohydride. This process is set forth in
U.S. Pat. Nos. 4,613,665; 4,810,784 to Larm and U.S. Pat. No. 5,049,403
to Larm et al.
[0133] Unlike rapamycin, heparin acts on circulating proteins in the blood
and heparin need only make contact with blood to be effective.
Accordingly, if used in conjunction with a medical device, such as a
stent, it would preferably be only on the side that comes into contact
with the blood. For example, if heparin were to be administered via a
stent, it would only have to be on the inner surface of the stent to be
effective.
[0134] In an exemplary embodiment of the invention, a stent may be
utilized in combination with rapamycin and heparin to treat vascular
disease. In this exemplary embodiment, the heparin is immobilized to the
inner surface of the stent so that it is in contact with the blood and
the rapamycin is immobilized to the outer surface of the stent so that it
is in contact with the surrounding tissue. FIG. 7 illustrates a
cross-section of a band 102 of the stent 100 illustrated in FIG. 1. As
illustrated, the band 102 is coated with heparin 108 on its inner surface
110 and with rapamycin 112 on its outer surface 114.
[0135] In an alternate exemplary embodiment, the stent may comprise a
heparin layer immobilized on its inner surface, and rapamycin and heparin
on its outer surface. Utilizing current coating techniques, heparin tends
to form a stronger bond with the surface it is immobilized to then does
rapamycin. Accordingly, it may be possible to first immobilize the
rapamycin to the outer surface of the stent and then immobilize a layer
of heparin to the rapamycin layer. In this embodiment, the rapamycin may
be more securely affixed to the stent while still effectively eluting
from its polymeric matrix, through the heparin and into the surrounding
tissue. FIG. 8 illustrates a cross-section of a band 102 of the stent 100
illustrated in FIG. 1. As illustrated, the band 102 is coated with
heparin 108 on its inner surface 110 and with rapamycin 112 and heparin
108 on its outer surface 114.
[0136] There are a number of possible ways to immobilize, i.e., entrapment
or covalent linkage with an erodible bond, the heparin layer to the
rapamycin layer. For example, heparin may be introduced into the top
layer of the polymeric matrix. In other embodiments, different forms of
heparin may be directly immobilized onto the top coat of the polymeric
matrix, for example, as illustrated in FIG. 9. As illustrated, a
hydrophobic heparin layer 116 may be immobilized onto the top coat layer
118 of the rapamycin layer 112. A hydrophobic form of heparin is utilized
because rapamycin and heparin coatings represent incompatible coating
application technologies. Rapamycin is an organic solvent-based coating
and heparin, in its native form, is a water-based coating.
[0137] As stated above, a rapamycin coating may be applied to stents by a
dip, spray or spin coating method, and/or any combination of these
methods. Various polymers may be utilized. For example, as described
above, poly(ethylene-co-vinyl acetate) and polybutyl methacrylate blends
may be utilized. Other polymers may also be utilized, but not limited to,
for example, polyvinylidene fluoride-co-hexafluoropropylene and
polyethylbutyl methacrylate-co-hexyl methacrylate. Also as described
above, barrier or top coatings may also be applied to modulate the
dissolution of rapamycin from the polymer matrix. In the exemplary
embodiment described above, a thin layer of heparin is applied to the
surface of the polymeric matrix. Because these polymer systems are
hydrophobic and incompatible with the hydrophilic heparin, appropriate
surface modifications may be required.
[0138] The application of heparin to the surface of the polymeric matrix
may be performed in various ways and utilizing various biocompatible
materials. For example, in one embodiment, in water or alcoholic
solutions, polyethylene imine may be applied on the stents, with care not
to degrade the rapamycin (e.g., pH<7, low temperature), followed by
the application of sodium heparinate in aqueous or alcoholic solutions.
As an extension of this surface modification, covalent heparin may be
linked on polyethylene imine using amide-type chemistry (using a
carbondiimide activator, e.g. EDC) or reductive amination chemistry
(using CBAS-heparin and sodium cyanoborohydride for coupling). In another
exemplary embodiment, heparin may be photolinked on the surface, if it is
appropriately grafted with photo initiator moieties. Upon application of
this modified heparin formulation on the covalent stent surface, light
exposure causes cross-linking and immobilization of the heparin on the
coating surface. In yet another exemplary embodiment, heparin may be
complexed with hydrophobic quaternary ammonium salts, rendering the
molecule soluble in organic solvents (e.g. benzalkonium heparinate,
troidodecylmethylammonium heparinate). Such a formulation of heparin may
be compatible with the hydrophobic rapamycin coating, and may be applied
directly on the coating surface, or in the rapamycin/hydrophobic polymer
formulation.
[0139] It is important to note that the stent, as described above, may be
formed from any number of materials, including various metals, polymeric
materials and ceramic materials. Accordingly, various technologies may be
utilized to immobilize the various drugs, agent, compound combinations
thereon. Specifically, in addition to the polymeric matricies, described
above, biopolymers may be utilized. Biopolymers may be generally
classified as natural polymers, while the above-described polymers may be
described as synthetic polymers. Exemplary biopolymers, which may be
utilized include agarose, alginate, gelatin, collagen and elastin. In
addition, the drugs, agents or compounds may be utilized in conjunction
with other percutaneously delivered medical devices such as grafts and
perfusion balloons.
[0140] In addition to utilizing an anti-proliferative and anti-coagulant,
antiinflammatories may also be utilized in combination therewith. One
example of such a combination would be the addition of an
anti-inflammatory corticosteroid such as dexamethasone with an
anti-proliferative, such as rapamycin, cladribine, vincristine, taxol, or
a nitric oxide donor and an anti-coagulant, such as heparin. Such
combination therapies might result in a better therapeutic effect, i.e.,
less proliferation as well as less inflammation, a stimulus for
proliferation, than would occur with either agent alone. The delivery of
a stent comprising an anti-proliferative, anti-coagulant, and an
anti-inflammatory to an injured vessel would provide the added
therapeutic benefit of limiting the degree of local smooth muscle cell
proliferation, reducing a stimulus for proliferation, i.e., inflammation
and reducing the effects of coagulation thus enhancing the
restenosis-limiting action of the stent.
[0141] In other exemplary embodiments of the inventions, growth factor
inhibitor or cytokine signal transduction inhibitor, such as the ras
inhibitor, R115777 or P38 kinase inhibitor RWJ67657, or a tyrosine kinase
inhibitor, such as tyrphostin, might be combined with an
anti-proliferative agent such as taxol, vincristine or rapamycin so that
proliferation of smooth muscle cells could be inhibited by different
mechanisms. Alternatively, an anti-proliferative agent such as taxol,
vincristine or rapamycin could be combined with an inhibitor of
extracellular matrix synthesis such as halofuginone. In the above cases,
agents acting by different mechanisms could act synergistically to reduce
smooth muscle cell proliferation and vascular hyperplasia. This invention
is also intended to cover other combinations of two or more such drug
agents. As mentioned above, such drugs, agents or compounds could be
administered systemically, delivered locally via drug delivery catheter,
or formulated for delivery from the surface of a stent, or given as a
combination of systemic and local therapy.
[0142] In addition to anti-proliferatives, anti-inflammatories and
anti-coagulants, other drugs, agents or compounds may be utilized in
conjunction with the medical devices. For example, immunosuppressants may
be utilized alone or in combination with these other drugs, agents or
compounds. Also gene therapy delivery mechanisms such as modified genes
(nucleic acids including recombinant DNA) in viral vectors and non-viral
gene vectors such as plasmids may also be introduced locally via a
medical device. In addition, the present invention may be utilized with
cell based therapy.
[0143] In addition to all of the drugs, agents, compounds and modified
genes described above, chemical agents that are not ordinarily
therapeutically or biologically active may also be utilized in
conjunction with the present invention. These chemical agents, commonly
referred to as pro-drugs, are agents that become biologically active upon
their introduction into the living organism by one or more mechanisms.
These mechanisms include the addition of compounds supplied by the
organism or the cleavage of compounds from the agents caused by another
agent supplied by the organism. Typically, pro-drugs are more absorbable
by the organism. In addition, pro-drugs may also provide some additional
measure of time release.
[0144] The coatings and drugs, agents or compounds described above may be
utilized in combination with any number of medical devices, and in
particular, with implantable medical devices such as stents and
stent-grafts. Other devices such as vena cava filters and anastomosis
devices may be used with coatings having drugs, agents or compounds
therein. The exemplary stent illustrated in FIGS. 1 and 2 is a balloon
expandable stent. Balloon expandable stents may be utilized in any number
of vessels or conduits, and are particularly well suited for use in
coronary arteries. Self-expanding stents, on the other hand, are
particularly well suited for use in vessels where crush recovery is a
critical factor, for example, in the carotid artery. Accordingly, it is
important to note that any of the drugs, agents or compounds, as well as
the coatings described above, may be utilized in combination with
self-expanding stents such as those described below.
[0145] There is illustrated in FIGS. 10 and 11, a stent 200, which may be
utilized in connection with the present invention. FIGS. 10 and 11
illustrate the exemplary stent 200 in its unexpanded or compressed state.
The stent 200 is preferably made from a superelastic alloy such as
Nitinol. Most preferably, the stent 200 is made from an alloy comprising
from about fifty percent (as used herein these percentages refer to
weight percentages) Ni to about sixty percent Ni, and more preferably
about 55.8 percent Ni, with the remainder of the alloy being Ti.
Preferably, the stent 200 is designed such that it is superelastic at
body temperature, and preferably has an Af in the range from about
twenty-four degrees C. to about thirty-seven degrees C. The superelastic
design of the stent 200 makes it crush recoverable which, as discussed
above, makes it useful as a stent or frame for any number of vascular
devices in different applications.
[0146] Stent 200 is a tubular member having front and back open ends 202
and 204 and a longitudinal axis 206 extending therebetween. The tubular
member has a first smaller diameter, FIGS. 10 and 11, for insertion into
a patient and navigation through the vessels, and a second larger
diameter, FIGS. 12 and 13, for deployment into the target area of a
vessel. The tubular member is made from a plurality of adjacent hoops
208, FIG. 10 showing hoops 208(a)-208(d), extending between the front and
back ends 202 and 204. The hoops 208 include a plurality of longitudinal
struts 210 and a plurality of loops 212 connecting adjacent struts,
wherein adjacent struts are connected at opposite ends so as to form a
substantially S or Z shape pattern. The loops 212 are curved,
substantially semi-circular with symmetrical sections about their centers
214.
[0147] Stent 200 further includes a plurality of bridges 216 which connect
adjacent hoops 208 and which can best be described in detail by referring
to FIG. 14. Each bridge 216 has two ends 218 and 220. The bridges 216
have one end attached to one strut and/or loop, and another end attached
to a strut and/or loop on an adjacent hoop. The bridges 216 connect
adjacent struts together at bridge to loop connection points 222 and 224.
For example, bridge end 218 is connected to loop 214(a) at bridge to loop
connection point 222, and bridge end 220 is connected to loop 214(b) at
bridge to loop connection point 224. Each bridge to loop connection point
has a center 226. The bridge to loop connection points are separated
angularly with respect to the longitudinal axis. That is, the connection
points are not immediately opposite each other. Essentially, one could
not draw a straight line between the connection points, wherein such line
would be parallel to the longitudinal axis of the stent.
[0148] The above described geometry helps to better distribute strain
throughout the stent, prevents metal to metal contact when the stent is
bent, and minimizes the opening size between the struts, loops and
bridges. The number of and nature of the design of the struts, loops and
bridges are important factors when determining the working properties and
fatigue life properties of the stent. It was previously thought that in
order to improve the rigidity of the stent, that struts should be large,
and therefore there should be fewer struts per hoop. However, it has now
been discovered that stents having smaller struts and more struts per
hoop actually improve the construction of the stent and provide greater
rigidity. Preferably, each hoop has between twenty-four to thirty-six or
more struts. It has been determined that a stent having a ratio of number
of struts per hoop to strut length L (in inches) which is greater than
four hundred has increased rigidity over prior art stents, which
typically have a ratio of under two hundred. The length of a strut is
measured in its compressed state parallel to the longitudinal axis 206 of
the stent 200 as illustrated in FIG. 10.
[0149] As seen from a comparison of FIGS. 10 and 12, the geometry of the
stent 200 changes quite significantly as the stent 200 is deployed from
its unexpanded state to its expanded state. As a stent undergoes
diametric change, the strut angle and strain levels in the loops and
bridges are affected. Preferably, all of the stent features will strain
in a predictable manor so that the stent is reliable and uniform in
strength. In addition, it is preferable to minimize the maximum strain
experienced by struts loops and bridges, since Nitinol properties are
more generally limited by strain rather than by stress. As will be
discussed in greater detail below, the stent sits in the delivery system
in its unexpanded state as shown in FIGS. 19 and 20. As the stent is
deployed, it is allowed to expand towards its expanded state, as shown in
FIG. 12, which preferably has a diameter which is the same or larger than
the diameter of the target vessel. Nitinol stents made from wire deploy
in much the same manner, and are dependent upon the same design
constraints, as laser cut stents. Stainless steel stents deploy similarly
in terms of geometric changes as they are assisted by forces from
balloons or other devices.
[0150] In trying to minimize the maximum strain experienced by features of
the stent, the present invention utilizes structural geometries which
distribute strain to areas of the stent which are less susceptible to
failure than others. For example, one of the most vulnerable areas of the
stent is the inside radius of the connecting loops. The connecting loops
undergo the most deformation of all the stent features. The inside radius
of the loop would normally be the area with the highest level of strain
on the stent. This area is also critical in that it is usually the
smallest radius on the stent. Stress concentrations are generally
controlled or minimized by maintaining the largest radii possible.
Similarly, we want to minimize local strain concentrations on the bridge
and bridge connection points. One way to accomplish this is to utilize
the largest possible radii while maintaining feature widths which are
consistent with applied forces. Another consideration is to minimize the
maximum open area of the stent. Efficient utilization of the original
tube from which the stent is cut increases stent strength and its ability
to trap embolic material.
[0151] Many of these design objectives have been accomplished by an
exemplary embodiment of the present invention, illustrated in FIGS. 10,
11 and 14. As seen from these figures, the most compact designs which
maintain the largest radii at the loop to bridge connections are
non-symmetric with respect to the centerline of the strut connecting
loop. That is, loop to bridge connection point centers 226 are offset
from the center 214 of the loops 212 to which they are attached. This
feature is particularly advantageous for stents having large expansion
ratios, which in turn requires them to have extreme bending requirements
where large elastic strains are required. Nitinol can withstand extremely
large amounts of elastic strain deformation, so the above features are
well suited to stents made from this alloy. This feature allows for
maximum utilization of Ni--Ti or other material properties to enhance
radial strength, to improve stent strength uniformity, to improve fatigue
life by minimizing local strain levels, to allow for smaller open areas
which enhance entrapment of embolic material, and to improve stent
apposition in irregular vessel wall shapes and curves.
[0152] As seen in FIG. 14, stent 200 comprises strut connecting loops 212
having a width WI, as measured at the center 214 parallel to axis 206,
which are greater than the strut widths W2, as measured perpendicular to
axis 206 itself. In fact, it is preferable that the thickness of the
loops vary so that they are thickest near their centers. This increases
strain deformation at the strut and reduces the maximum strain levels at
the extreme radii of the loop. This reduces the risk of stent failure and
allows one to maximize radial strength properties. This feature is
particularly advantageous for stents having large expansion ratios, which
in turn requires them to have extreme bending requirements where large
elastic strains are required. Nitinol can withstand extremely large
amounts of elastic strain deformation, so the above features are well
suited to stents made from this alloy. As stated above, this feature
allows for maximum utilization of Ni--Ti or other material properties to
enhance radial strength, to improve stent strength uniformity, to improve
fatigue life by minimizing local strain levels, to allow for smaller open
areas which enhance entrapment of embolic material, and to improve stent
apposition in irregular vessel wall shapes and curves.
[0153] As mentioned above, bridge geometry changes as a stent is deployed
from its compressed state to its expanded state and vise-versa. As a
stent undergoes diametric change, strut angle and loop strain is
affected. Since the bridges are connected to either the loops, struts or
both, they are affected. Twisting of one end of the stent with respect to
the other, while loaded in the stent delivery system, should be avoided.
Local torque delivered to the bridge ends displaces the bridge geometry.
If the bridge design is duplicated around the stent perimeter, this
displacement causes rotational shifting of the two loops being connected
by the bridges. If the bridge design is duplicated throughout the stent,
as in the present invention, this shift will occur down the length of the
stent. This is a cumulative effect as one considers rotation of one end
with respect to the other upon deployment. A stent delivery system, such
as the one described below, will deploy the distal end first, then allow
the proximal end to expand. It would be undesirable to allow the distal
end to anchor into the vessel wall while holding the stent fixed in
rotation, then release the proximal end. This could cause the stent to
twist or whip in rotation to equilibrium after it is at least partially
deployed within the vessel. Such whipping action may cause damage to the
vessel.
[0154] However, one exemplary embodiment of the present invention, as
illustrated in FIGS. 10 and 11, reduces the chance of such events
happening when deploying the stent. By mirroring the bridge geometry
longitudinally down the stent, the rotational shift of the Z-sections or
S-sections may be made to alternate and will minimize large rotational
changes between any two points on a given stent during deployment or
constraint. That is, the bridges 216 connecting loop 208(b) to loop
208(c) are angled upwardly from left to right, while the bridges
connecting loop 208(c) to loop 208(d) are angled downwardly from left to
right. This alternating pattern is repeated down the length of the stent
200. This alternating pattern of bridge slopes improves the torsional
characteristics of the stent so as to minimize any twisting or rotation
of the stent with respect to any two hoops. This alternating bridge slope
is particularly beneficial if the stent starts to twist in vivo. As the
stent twists, the diameter of the stent will change. Alternating bridge
slopes tend to minimize this effect. The diameter of a stent having
bridges which are all sloped in the same direction will tend to grow if
twisted in one direction and shrink if twisted in the other direction.
With alternating bridge slopes this effect is minimized and localized.
[0155] The feature is particularly advantageous for stents having large
expansion ratios, which in turn requires them to have extreme bending
requirements where large elastic strains are required. Nitinol, as stated
above, can withstand extremely large amounts of elastic strain
deformation, so the above features are well suited to stents made from
this alloy. This feature allows for maximum utilization of Ni--Ti or
other material properties to enhance radial strength, to improve stent
strength uniformity, to improve fatigue life by minimizing local strain
levels, to allow for smaller open areas which enhance entrapment of
embolic material, and to improve stent apposition in irregular vessel
wall shapes and curves.
[0156] Preferably, stents are laser cut from small diameter tubing. For
prior art stents, this manufacturing process led to designs with
geometric features, such as struts, loops and bridges, having axial
widths W2, W1 and W3, respectively, which are larger than the tube wall
thickness T (illustrated in FIG. 12). When the stent is compressed, most
of the bending occurs in the plane that is created if one were to cut
longitudinally down the stent and flatten it out. However, for the
individual bridges, loops and struts, which have widths greater than
their thickness, there is a greater resistance to this in-plane bending
than to out-of-plane bending. Because of this, the bridges and struts
tend to twist, so that the stent as a whole may bend more easily. This
twisting is a buckling condition which is unpredictable and can cause
potentially high strain.
[0157] However, this problem has been solved in an exemplary embodiment of
the present invention, as illustrated in FIGS. 10-14. As seen from these
figures, the widths of the struts, hoops and bridges are equal to or less
than the wall thickness of the tube. Therefore, substantially all bending
and, therefore, all strains are "out-of-plane." This minimizes twisting
of the stent which minimizes or eliminates buckling and unpredictable
strain conditions. This feature is particularly advantageous for stents
having large expansion ratios, which in turn requires them to have
extreme bending requirements where large elastic strains are required.
Nitinol, as stated above, can withstand extremely large amounts of
elastic strain deformation, so the above features are well suited to
stents made from this alloy. This feature allows for maximum utilization
of Ni--Ti or other material properties to enhance radial strength, to
improve stent strength uniformity, to improve fatigue life by minimizing
local strain levels, to allow for smaller open areas which enhance
entrapment of embolic material, and to improve stent apposition in
irregular vessel wall shapes and curves.
[0158] An alternate exemplary embodiment of a stent that may be utilized
in conjunction with the present invention is illustrated in FIG. 15. FIG.
15 shows stent 300 which is similar to stent 200 illustrated in FIGS.
10-14. Stent 300 is made from a plurality of adjacent hoops 302, FIG. 15
showing hoops 302(a)-302(d). The hoops 302 include a plurality of
longitudinal struts 304 and a plurality of loops 306 connecting adjacent
struts, wherein adjacent struts are connected at opposite ends so as to
form a substantially S or Z shape pattern. Stent 300 further includes a
plurality of bridges 308 which connect adjacent hoops 302. As seen from
the figure, bridges 308 are non-linear and curve between adjacent hoops.
Having curved bridges allows the bridges to curve around the loops and
struts so that the hoops can be placed closer together which in turn,
minimizes the maximum open area of the stent and increases its radial
strength as well. This can best be explained by referring to FIG. 13. The
above described stent geometry attempts to minimize the largest circle
which could be inscribed between the bridges, loops and struts, when the
stent is expanded. Minimizing the size of this theoretical circle,
greatly improves the stent because it is then better suited to trap
embolic material once it is inserted into the patient.
[0159] As mentioned above, it is preferred that the stent of the present
invention be made from a superelastic alloy and most preferably made of
an alloy material having greater than 50.5 atomic percentage Nickel and
the balance Titanium. Greater than 50.5 atomic percentage Nickel allows
for an alloy in which the temperature at which the martensite phase
transforms completely to the austenite phase (the Af temperature) is
below human body temperature, and preferably is about twenty-four degrees
C. to about thirty-seven degrees C., so that austenite is the only stable
phase at body temperature.
[0160] In manufacturing the Nitinol stent, the material is first in the
form of a tube. Nitinol tubing is commercially available from a number of
suppliers including Nitinol Devices and Components, Fremont CA. The
tubular member is then loaded into a machine which will cut the
predetermined pattern of the stent into the tube, as discussed above and
as shown in the figures. Machines for cutting patterns in tubular devices
to make stents or the like are well known to those of ordinary skill in
the art and are commercially available. Such machines typically hold the
metal tube between the open ends while a cutting laser, preferably under
microprocessor control, cuts the pattern. The pattern dimensions and
styles, laser positioning requirements, and other information are
programmed into a microprocessor which controls all aspects of the
process. After the stent pattern is cut, the stent is treated and
polished using any number of methods or combination of methods well known
to those skilled in the art. Lastly, the stent is then cooled until it is
completely martensitic, crimped down to its un-expanded diameter and then
loaded into the sheath of the delivery apparatus.
[0161] As stated in previous sections of this application, markers having
a radiopacity greater than that of the superelastic alloys may be
utilized to facilitate more precise placement of the stent within the
vasculature. In addition, markers may be utilized to determine when and
if a stent is fully deployed. For example, by determining the spacing
between the markers, one can determine if the deployed stent has achieved
its maximum diameter and adjusted accordingly utilizing a tacking
process. FIG. 16 illustrates an exemplary embodiment of the stent 200
illustrated in FIGS. 10-14 having at least one marker on each end
thereof. In a preferred embodiment, a stent having thirty-six struts per
hoop can accommodate six markers 800. Each marker 800 comprises a marker
housing 802 and a marker insert 804. The marker insert 804 may be formed
from any suitable biocompatible material having a high radiopacity under
X-ray fluoroscopy. In other words, the marker inserts 804 should
preferably have a radiopacity higher than that of the material comprising
the stent 200. The addition of the marker housings 802 to the stent
necessitates that the lengths of the struts in the last two hoops at each
end of the stent 200 be longer than the strut lengths in the body of the
stent to increase the fatigue life at the stent ends. The marker housings
802 are preferably cut from the same tube as the stent as briefly
described above. Accordingly, the housings 802 are integral to the stent
200. Having the housings 802 integral to the stent 200 serves to ensure
that the markers 800 do not interfere with the operation of the stent.
[0162] FIG. 17 is a cross-sectional view of a marker housing 802. The
housing 802 may be elliptical when observed from the outer surface as
illustrated in FIG. 16. As a result of the laser cutting process, the
hole 806 in the marker housing 802 is conical in the radial direction
with the outer surface 808 having a diameter larger than the diameter of
the inner surface 810, as illustrated in FIG. 17. The conical tapering in
the marker housing 802 is beneficial in providing an interference fit
between the marker insert 804 and the marker housing 802 to prevent the
marker insert 804 from being dislodged once the stent 200 is deployed. A
detailed description of the process of locking the marker insert 804 into
the marker housing 802 is given below.
[0163] As set forth above, the marker inserts 804 may be made from any
suitable material having a radiopacity higher than the superelastic
material forming the stent or other medical device. For example, the
marker insert 804 may be formed from niobium, tungsten, gold, platinum or
tantalum. In the preferred embodiment, tantalum is utilized because of
its closeness to nickel-titanium in the galvanic series and thus would
minimize galvanic corrosion. In addition, the surface area ratio of the
tantalum marker inserts 804 to the nickel-titanium is optimized to
provide the largest possible tantalum marker insert, easy to see, while
minimizing the galvanic corrosion potential. For example, it has been
determined that up to nine marker inserts 804 having a diameter of 0.010
inches could be placed at the end of the stent 200; however, these marker
inserts 804 would be less visible under X-ray fluoroscopy. On the other
hand, three to four marker inserts 804 having a diameter of 0.025 inches
could be accommodated on the stent 200; however, galvanic corrosion
resistance would be compromised. Accordingly, in the preferred
embodiment, six tantalum markers having a diameter of 0.020 inches are
utilized on each end of the stent 200 for a total of twelve markers 800.
[0164] The tantalum markers 804 may be manufactured and loaded into the
housing utilizing a variety of known techniques. In the exemplary
embodiment, the tantalum markers 804 are punched out from an annealed
ribbon stock and are shaped to have the same curvature as the radius of
the marker housing 802 as illustrated in FIG. 17. Once the tantalum
marker insert 804 is loaded into the marker housing 802, a coining
process is used to properly seat the marker insert 804 below the surface
of the housing 802. The coining punch is also shaped to maintain the same
radius of curvature as the marker housing 802. As illustrated in FIG. 17,
the coining process deforms the marker housing 802 material to lock in
the marker insert 804.
[0165] As stated above, the hole 806 in the marker housing 802 is conical
in the radial direction with the outer surface 808 having a diameter
larger than the diameter of the inner surface 810 as illustrated in FIG.
17. The inside and outside diameters vary depending on the radius of the
tube from which the stent is cut. The marker inserts 804, as stated
above, are formed by punching a tantalum disk from annealed ribbon stock
and shaping it to have the same radius of curvature as the marker housing
802. It is important to note that the marker inserts 804, prior to
positioning in the marker housing 804, have straight edges. In other
words, they are not angled to match the hole 806. The diameter of the
marker insert 804 is between the inside and outside diameter of the
marker housing 802. Once the marker insert 804 is loaded into the marker
housing, a coining process is used to properly seat the marker insert 804
below the surface of the housing 802. In the preferred embodiment, the
thickness of the marker insert 804 is less than or equal to the thickness
of the tubing and thus the thickness or height of the hole 806.
Accordingly, by applying the proper pressure during the coining process
and using a coining tool that is larger than the marker insert 804, the
marker insert 804 may be seated in the marker housing 802 in such a way
that it is locked into position by a radially oriented protrusion 812.
Essentially, the applied pressure, and size and shape of the housing tool
forces the marker insert 804 to form the protrusion 812 in the marker
housing 802. The coining tool is also shaped to maintain the same radius
of curvature as the marker housing. As illustrated in FIG. 17, the
protrusion 812 prevents the marker insert 804 from becoming dislodged
from the marker housing.
[0166] It is important to note that the marker inserts 804 are positioned
in and locked into the marker housing 802 when the stent 200 is in its
unexpanded state. This is due to the fact that it is desirable that the
tube's natural curvature be utilized. If the stent were in its expanded
state, the coining process would change the curvature due to the pressure
or force exerted by the coining tool.
[0167] As illustrated in FIG. 18, the marker inserts 804 form a
substantially solid line that clearly defines the ends of the stent in
the stent delivery system when seen under fluoroscopic equipment. As the
stent 200 is deployed from the stent delivery system, the markers 800
move away from each other and flower open as the stent 200 expands as
illustrated in FIG. 16. The change in the marker grouping provides the
physician or other health care provider with the ability to determine
when the stent 200 has been fully deployed from the stent delivery
system.
[0168] It is important to note that the markers 800 may be positioned at
other locations on the stent 200.
[0169] It is believed that many of the advantages of the present invention
can be better understood through a brief description of a delivery
apparatus for the stent, as shown in FIGS. 19 and 20. FIGS. 19 and 20
show a self-expanding stent delivery apparatus 10 for a stent made in
accordance with the present invention. Apparatus 10 comprises inner and
outer coaxial tubes. The inner tube is called the shaft 12 and the outer
tube is called the sheath 14. Shaft 12 has proximal and distal ends. The
proximal end of the shaft 12 terminates at a luer lock hub 16.
Preferably, shaft 12 has a proximal portion 18 which is made from a
relatively stiff material such as stainless steel, Nitinol, or any other
suitable material, and a distal portion 20 which may be made from a
polyethylene, polyimide, Pellethane, Pebax, Vestamid, Cristamid,
Grillamid or any other suitable material known to those of ordinary skill
in the art. The two portions are joined together by any number of means
known to those of ordinary skill in the art. The stainless steel proximal
end gives the shaft the necessary rigidity or stiffness it needs to
effectively push out the stent, while the polymeric distal portion
provides the necessary flexibility to navigate tortuous vessels.
[0170] The distal portion 20 of the shaft 12 has a distal tip 22 attached
thereto. The distal tip 22 has a proximal end 24 whose diameter is
substantially the same as the outer diameter of the sheath 14. The distal
tip 22 tapers to a smaller diameter from its proximal end to its distal
end, wherein the distal end 26 of the distal tip 22 has a diameter
smaller than the inner diameter of the sheath 14. Also attached to the
distal portion 20 of shaft 12 is a stop 28 which is proximal to the
distal tip 22. Stop 28 may be made from any number of materials known in
the art, including stainless steel, and is even more preferably made from
a highly radiopaque material such as platinum, gold or tantalum. The
diameter of stop 28 is substantially the same as the inner diameter of
sheath 14, and would actually make frictional contact with the inner
surface of the sheath. Stop 28 helps to push the stent out of the sheath
during deployment, and helps keep the stent from migrating proximally
into the sheath 14.
[0171] A stent bed 30 is defined as being that portion of the shaft
between the distal tip 22 and the stop 28. The stent bed 30 and the stent
200 are coaxial so that the distal portion 20 of shaft 12 comprising the
stent bed 30 is located within the lumen of the stent 200. However, the
stent bed 30 does not make any contact with stent 200 itself. Lastly,
shaft 12 has a guidewire lumen 32 extending along its length from its
proximal end and exiting through its distal tip 22. This allows the shaft
12 to receive a guidewire much in the same way that an ordinary balloon
angioplasty catheter receives a guidewire. Such guidewires are well known
in art and help guide catheters and other medical devices through the
vasculature of the body.
[0172] Sheath 14 is preferably a polymeric catheter and has a proximal end
terminating at a sheath hub 40. Sheath 14 also has a distal end which
terminates at the proximal end 24 of distal tip 22 of the shaft 12, when
the stent is in its fully un-deployed position as shown in the figures.
The distal end of sheath 14 includes a radiopaque marker band 34 disposed
along its outer surface. As will be explained below, the stent is fully
deployed from the delivery apparatus when the marker band 34 is lined up
with radiopaque stop 28, thus indicating to the physician that it is now
safe to remove the apparatus 10 from the body. Sheath 14 preferably
comprises an outer polymeric layer and an inner polymeric layer.
Positioned between outer and inner layers is a braided reinforcing layer.
Braided reinforcing layer is preferably made from stainless steel. The
use of braided reinforcing layers in other types of medical devices can
be found in U.S. Pat. No. 3,585,707 issued to Stevens on Jun. 22, 1971,
U.S. Pat. No. 5,045,072 issued to Castillo et al. on Sep. 3, 1991, and
U.S. Pat. No. 5,254,107 issued to Soltesz on Oct. 19, 1993.
[0173] FIGS. 19 and 20 illustrate the stent 200 as being in its fully
un-deployed position. This is the position the stent is in when the
apparatus 10 is inserted into the vasculature and its distal end is
navigated to a target site. Stent 200 is disposed around stent bed 30 and
at the distal end of sheath 14. The distal tip 22 of the shaft 12 is
distal to the distal end of the sheath 14, and the proximal end of the
shaft 12 is proximal to the proximal end of the sheath 14. The stent 200
is in a compressed state and makes frictional contact with the inner
surface 36 of the sheath 14.
[0174] When being inserted into a patient, sheath 14 and shaft 12 are
locked together at their proximal ends by a Tuohy Borst valve 38. This
prevents any sliding movement between the shaft and sheath which could
result in a premature deployment or partial deployment of the stent 200.
When the stent 200 reaches its target site and is ready for deployment,
the Tuohy Borst valve 38 is opened so that that the sheath 14 and shaft
12 are no longer locked together.
[0175] The method under which the apparatus 10 deploys the stent 200 is
readily apparent. The apparatus 10 is first inserted into the vessel
until the radiopaque stent markers 800 (front 202 and back 204 ends, see
FIG. 16) are proximal and distal to the target lesion. Once this has
occurred the physician would open the Tuohy Borst valve 38. The physician
would then grasp hub 16 of shaft 12 so as to hold it in place.
Thereafter, the physician would grasp the proximal end of the sheath 14
and slide it proximal, relative to the shaft 12. Stop 28 prevents the
stent 200 from sliding back with the sheath 14, so that as the sheath 14
is moved back, the stent 200 is pushed out of the distal end of the
sheath 14. As stent 200 is being deployed the radiopaque stent markers
800 move apart once they come out of the distal end of sheath 14. Stent
deployment is complete when the marker 34 on the outer sheath 14 passes
the stop 28 on the inner shaft 12. The apparatus 10 can now be withdrawn
through the stent 200 and removed from the patient.
[0176] FIG. 21 illustrates the stent 200 in a partially deployed state. As
illustrated, as the stent 200 expands from the delivery device 10, the
markers 800 move apart from one another and expand in a flower like
manner.
[0177] It is important to note that any of the above-described medical
devices may be coated with coatings that comprise drugs, agents or
compounds or simply with coatings that contain no drugs, agents or
compounds. In addition, the entire medical device may be coated or only a
portion of the device may be coated. The coating may be uniform or
non-uniform. The coating may be discontinuous. However, the markers on
the stent are preferably coated in a manner so as to prevent coating
buildup which may interfere with the operation of the device.
[0178] In a preferred exemplary embodiment, the self-expanding stents,
described above, may be coated with a rapamycin containing polymer. In
this embodiment, the polymeric coated stent comprises rapamycin in an
amount ranging from about fifty to one-thousand micrograms per square
centimeter surface area of the vessel that is spanned by the stent. The
rapamycin is mixed with the polyvinylidenefluoride-hexafluoropropylene
polymer (described above) in the ratio of drug to polymer of about
thirty/seventy. The polymer is made by a batch process using the two
monomers, vinylidene fluoride and hexafluoropropylene under high pressure
by an emulsion polymerization process. In an alternate exemplary
embodiment, the polymer may be made utilizing a batch dispersion process.
The polymeric coating weight itself is in the range from about
two-hundred to about one thousand seven hundred micrograms per square
centimeter surface area of the vessel that is spanned by the stent.
[0179] The coated stent comprises a base coat, commonly referred to as a
primer layer. The primer layer typically improves the adhesion of the
coating layer that comprises the rapamycin. The primer also facilitates
uniform wetting of the surface thereby enabling the production of a
uniform rapamycin containing coating. The primer layer may be applied
using any of the above-described techniques. It is preferably applied
utilizing a dip coating process. The primer coating is in the range from
about one to about ten percent of the total weight of the coating. The
next layer applied is the rapamycin containing layer. The rapamycin
containing layer is applied by a spin coating process and subsequently
dried in a vacuum oven for approximately sixteen hours at a temperature
in the range from about fifty to sixty degrees centigrade. After drying
or curing, the stent is mounted onto a stent delivery catheter using a
process similar to the uncoated stent. The mounted stent is then packaged
and sterilized in any number of ways. In one exemplary embodiment, the
stent is sterilized using ethylene oxide.
[0180] As described above, various drugs, agents or compounds may be
locally delivered via medical devices. For example, rapamycin and heparin
may be delivered by a stent to reduce restenosis, inflammation, and
coagulation. Various techniques for immobilizing the drugs, agents or
compounds are discussed above, however, maintaining the drugs, agents or
compounds on the medical devices during delivery and positioning is
critical to the success of the procedure or treatment. For example,
removal of the drug, agent or compound coating during delivery of the
stent can potentially cause failure of the device. For a self-expanding
stent, the retraction of the restraining sheath may cause the drugs,
agents or compounds to rub off the stent. For a balloon expandable stent,
the expansion of the balloon may cause the drugs, agents or compounds to
simply delaminate from the stent through contact with the balloon or via
expansion. Therefore, prevention of this potential problem is important
to have a successful therapeutic medical device, such as a stent.
[0181] There are a number of approaches that may be utilized to
substantially reduce the above-described concern. In one exemplary
embodiment, a lubricant or mold release agent may be utilized. The
lubricant or mold release agent may comprise any suitable biocompatible
lubricious coating. An exemplary lubricious coating may comprise
silicone. In this exemplary embodiment, a solution of the silicone base
coating may be introduced onto the balloon surface, onto the polymeric
matrix, and/or onto the inner surface of the sheath of a self-expanding
stent delivery apparatus and allowed to air cure. Alternately, the
silicone based coating may be incorporated into the polymeric matrix. It
is important to note, however, that any number of lubricious materials
may be utilized, with the basic requirements being that the material be
biocompatible, that the material not interfere with the
actions/effectiveness of the drugs, agents or compounds and that the
material not interfere with the materials utilized to immobilize the
drugs, agents or compounds on the medical device. It is also important to
note that one or more, or all of the above-described approaches may be
utilized in combination.
[0182] Referring now to FIG. 22, there is illustrated a balloon 400 of a
balloon catheter that may be utilized to expand a stent in situ. As
illustrated, the balloon 400 comprises a lubricious coating 402. The
lubricious coating 402 functions to minimize or substantially eliminate
the adhesion between the balloon 400 and the coating on the medical
device. In the exemplary embodiment described above, the lubricious
coating 402 would minimize or substantially eliminate the adhesion
between the balloon 400 and the heparin or rapamycin coating. The
lubricious coating 402 may be attached to and maintained on the balloon
400 in any number of ways including but not limited to dipping, spraying,
brushing or spin coating of the coating material from a solution or
suspension followed by curing or solvent removal step as needed.
[0183] Materials such as synthetic waxes, e.g. diethyleneglycol
monostearate, hydrogenated castor oil, oleic acid, stearic acid, zinc
stearate, calcium stearate, ethylenebis (stearamide), natural products
such as paraffin wax, spermaceti wax, carnuba wax, sodium alginate,
ascorbic acid and flour, fluorinated compounds such as perfluoroalkanes,
perfluorofatty acids and alcohol, synthetic polymers such as silicones
e.g. polydimethylsiloxane, polytetrafluoroethylene, polyfluoroethers,
polyalkylglycol e.g. polyethylene glycol waxes, and inorganic materials
such as talc, kaolin, mica, and silica may be used to prepare these
coatings. Vapor deposition polymerization e.g. parylene-C deposition, or
RF-plasma polymerization of perfluoroalkenes and perfluoroalkanes can
also be used to prepare these lubricious coatings.
[0184] FIG. 23 illustrates a cross-section of a band 102 of the stent 100
illustrated in FIG. 1. In this exemplary embodiment, the lubricious
coating 500 is immobilized onto the outer surface of the polymeric
coating. As described above, the drugs, agents or compounds may be
incorporated into a polymeric matrix. The stent band 102 illustrated in
FIG. 23 comprises a base coat 502 comprising a polymer and rapamycin and
a top coat 504 or diffusion layer 504 also comprising a polymer. The
lubricious coating 500 is affixed to the top coat 502 by any suitable
means, including but not limited to spraying, brushing, dipping or spin
coating of the coating material from a solution or suspension with or
without the polymers used to create the top coat, followed by curing or
solvent removal step as needed. Vapor deposition polymerization and
RF-plasma polymerization may also be used to affix those lubricious
coating materials that lend themselves to this deposition method, to the
top coating. In an alternate exemplary embodiment, the lubricious coating
may be directly incorporated into the polymeric matrix.
[0185] If a self-expanding stent is utilized, the lubricious coating may
be affixed to the inner surface of the restraining sheath. FIG. 24
illustrates a self-expanding stent 200 (FIG. 10) within the lumen of a
delivery apparatus sheath 14. As illustrated, a lubricious coating 600 is
affixed to the inner surfaces of the sheath 14. Accordingly, upon
deployment of the stent 200, the lubricious coating 600 preferably
minimizes or substantially eliminates the adhesion between the sheath 14
and the drug, agent or compound coated stent 200.
[0186] In an alternate approach, physical and/or chemical cross-linking
methods may be applied to improve the bond strength between the polymeric
coating containing the drugs, agents or compounds and the surface of the
medical device or between the polymeric coating containing the drugs,
agents or compounds and a primer. Alternately, other primers applied by
either traditional coating methods such as dip, spray or spin coating, or
by RF-plasma polymerization may also be used to improve bond strength.
For example, as shown in FIG. 25, the bond strength can be improved by
first depositing a primer layer 700 such as vapor polymerized parylene-C
on the device surface, and then placing a second layer 702 which
comprises a polymer that is similar in chemical composition to the one or
more of the polymers that make up the drug-containing matrix 704, e.g.,
polyethylene-co-vinyl acetate or polybutyl methacrylate but has been
modified to contain cross-linking moieties. This secondary layer 702 is
then cross-linked to the primer after exposure to ultraviolet light. It
should be noted that anyone familiar with the art would recognize that a
similar outcome could be achieved using cross-linking agents that are
activated by heat with or without the presence of an activating agent.
The drug-containing matrix 704 is then layered onto the secondary layer
702 using a solvent that swells, in part or wholly, the secondary layer
702. This promotes the entrainment of polymer chains from the matrix into
the secondary layer 702 and conversely from the secondary layer 702 into
the drug-containing matrix 704. Upon removal of the solvent from the
coated layers, an interpenetrating or interlocking network of the polymer
chains is formed between the layers thereby increasing the adhesion
strength between them. A top coat 706 is used as described above.
[0187] A related difficulty occurs in medical devices such as stents. In
the drug-coated stents crimped state, some struts come into contact with
each other and when the stent is expanded, the motion causes the
polymeric coating comprising the drugs, agents or compounds to stick and
stretch. This action may potentially cause the coating to separate from
the stent in certain areas. The predominant mechanism of the coating
self-adhesion is believed to be due to mechanical forces. When the
polymer comes in contact with itself, its chains can tangle causing the
mechanical bond, similar to hook and loop fasteners such as Velcro.RTM..
Certain polymers do not bond with each other, for example,
fluoropolymers. For other polymers, however, powders may be utilized. In
other words, a powder may be applied to the one or more polymers
incorporating the drugs, agents or other compounds on the surfaces of the
medical device to reduce the mechanical bond. Any suitable biocompatible
material which does not interfere with the drugs, agents, compounds or
materials utilized to immobilize the drugs, agents or compounds onto the
medical device may be utilized. For example, a dusting with a water
soluble powder may reduce the tackiness of the coatings surface and this
will prevent the polymer from sticking to itself thereby reducing the
potential for delamination. The powder should be water-soluble so that it
does not present an emboli risk. The powder may comprise an anti-oxidant,
such as vitamin C, or it may comprise an anti-coagulant, such as aspirin
or heparin. An advantage of utilizing an anti-oxidant may be in the fact
that the anti-oxidant may preserve the other drugs, agents or compounds
over longer periods of time.
[0188] It is important to note that crystalline polymers are generally not
sticky or tacky. Accordingly, if crystalline polymers are utilized rather
than amorphous polymers, then additional materials may not be necessary.
It is also important to note that polymeric coatings without drugs,
agents, and/or compounds may improve the operating characteristics of the
medical device. For example, the mechanical properties of the medical
device may be improved by a polymeric coating, with or without drugs,
agents and/or compounds. A coated stent may have improved flexibility and
increased durability. In addition, the polymeric coating may
substantially reduce or eliminate galvanic corrosion between the
different metals comprising the medical device.
[0189] Any of the above-described medical devices may be utilized for the
local delivery of drugs, agents and/or compounds to other areas, not
immediately around the device itself. In order to avoid the potential
complications associated with systemic drug delivery, the medical devices
of the present invention may be utilized to deliver therapeutic agents to
areas adjacent to the medical device. For example, a rapamycin coated
stent may deliver the rapamycin to the tissues surrounding the stent as
well as areas upstream of the stent and downstream of the stent. The
degree of tissue penetration depends on a number of factors, including
the drug, agent or compound, the concentrations of the drug and the
release rate of the agent.
[0190] The drug, agent and/or compound/carrier or vehicle compositions
described above may be formulated in a number of ways. For example, they
may be formulated utilizing additional components or constituents,
including a variety of excipient agents and/or formulary components to
affect manufacturability, coating integrity, sterilizability, drug
stability, and drug release rate. Within exemplary embodiments of the
present invention, excipient agents and/or formulary components may be
added to achieve both fast-release and sustained-release drug elution
profiles. Such excipient agents may include salts and/or inorganic
compounds such as acids/bases or buffer components, anti-oxidants,
surfactants, polypeptides, proteins, carbohydrates including sucrose,
glucose or dextrose, chelating agents such as EDTA, glutathione or other
excipients or agents.
[0191] Although shown and described is what is believed to be the most
practical and preferred embodiments, it is apparent that departures from
specific designs and methods described and shown will suggest themselves
to those skilled in the art and may be used without departing from the
spirit and scope of the invention. The present invention is not
restricted to the particular constructions described and illustrated, but
should be constructed to cohere with all modifications that may fall
within the scope of the appended claims.
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